Low-field magnetic resonance imaging methods and apparatus

ABSTRACT

According to some aspects, a portable magnetic resonance imaging system is provided. The portable magnetic resonance imaging system comprises a B 0  magnet configured to produce a B 0  magnetic field for an imaging region of the magnetic resonance imaging system, a noise reduction system configured to detect and suppress at least some electromagnetic noise in an operating environment of the portable magnetic resonance imaging system, and electromagnetic shielding provided to attenuate at least some of the electromagnetic noise in the operating environment of the portable magnetic resonance imaging system, the electromagnetic shielding arranged to shield a fraction of the imaging region of the portable magnetic resonance imaging system. According to some aspects, the electromagnetic shielding may be configurable to providing a variable amount of shielding for the imaging region.

CROSS REFERENCE TO RELATED APPLICATIONS

This application claims priority under 35 U.S.C. § 119 to U.S. Provisional Application Ser. No. 62/425,465, filed Nov. 22, 2016, and titled LOW-FIELD MAGNETIC RESONANCE IMAGING METHODS AND APPARATUS, which is herein incorporated by reference in its entirety.

BACKGROUND

Magnetic resonance imaging (MRI) provides an important imaging modality for numerous applications and is widely utilized in clinical and research settings to produce images of the inside of the human body. As a generality, MRI is based on detecting magnetic resonance (MR) signals, which are electromagnetic waves emitted by atoms in response to state changes resulting from applied electromagnetic fields. For example, nuclear magnetic resonance (NMR) techniques involve detecting MR signals emitted from the nuclei of excited atoms upon the re-alignment or relaxation of the nuclear spin of atoms in an object being imaged (e.g., atoms in the tissue of the human body). Detected MR signals may be processed to produce images, which in the context of medical applications, allows for the investigation of internal structures and/or biological processes within the body for diagnostic, therapeutic and/or research purposes.

MRI provides an attractive imaging modality for biological imaging due to the ability to produce non-invasive images having relatively high resolution and contrast without the safety concerns of other modalities (e.g., without needing to expose the subject to ionizing radiation, e.g., x-rays, or introducing radioactive material to the body). Additionally, MRI is particularly well suited to provide soft tissue contrast, which can be exploited to image subject matter that other imaging modalities are incapable of satisfactorily imaging. Moreover, MR techniques are capable of capturing information about structures and/or biological processes that other modalities are incapable of acquiring. However, there are a number of drawbacks to MRI that, for a given imaging application, may involve the relatively high cost of the equipment, limited availability and/or difficulty in gaining access to clinical MRI scanners and/or the length of the image acquisition process.

The trend in clinical MRI has been to increase the field strength of MRI scanners to improve one or more of scan time, image resolution, and image contrast, which, in turn, continues to drive up costs. The vast majority of installed MRI scanners operate at 1.5 or 3 tesla (T), which refers to the field strength of the main magnetic field B₀. A rough cost estimate for a clinical MRI scanner is approximately one million dollars per tesla, which does not factor in the substantial operation, service, and maintenance costs involved in operating such MRI scanners.

Additionally, conventional high-field MRI systems typically require large superconducting magnets and associated electronics to generate a strong uniform static magnetic field (B₀) in which an object (e.g., a patient) is imaged. The size of such systems is considerable with a typical MRI installment including multiple rooms for the magnet, electronics, thermal management system, and control console areas. The size and expense of MRI systems generally limits their usage to facilities, such as hospitals and academic research centers, which have sufficient space and resources to purchase and maintain them. The high cost and substantial space requirements of high-field MRI systems results in limited availability of MRI scanners. As such, there are frequently clinical situations in which an MRI scan would be beneficial, but due to one or more of the limitations discussed above, is not practical or is impossible, as discussed in further detail below.

SUMMARY

Some embodiments include a portable magnetic resonance imaging system comprising a B₀ magnet configured to produce a B₀ magnetic field for an imaging region of the magnetic resonance imaging system, a noise reduction system configured to detect and suppress at least some electromagnetic noise in an operating environment of the portable magnetic resonance imaging system, and electromagnetic shielding provided to attenuate at least some of the electromagnetic noise in the operating environment of the portable magnetic resonance imaging system, the electromagnetic shielding arranged to shield a fraction of the imaging region of the portable magnetic resonance imaging system.

Some embodiments include a portable magnetic resonance imaging system comprising a B₀ magnet configured to produce a B₀ field for an imaging region of the magnetic resonance imaging system, a noise reduction system configured to detect and suppress at least some electromagnetic noise in an operating environment of the portable magnetic resonance imaging system, and electromagnetic shielding for at least a portion of the portable magnetic resonance imaging system, the electromagnetic shielding providing substantially no shielding of the imaging region of the portable magnetic resonance imaging system.

Some embodiments include portable magnetic resonance imaging system comprising a B₀ magnet configured to produce a B₀ field for an imaging region of the magnetic resonance imaging system, a housing for the B₀ magnet, and at least one electromagnetic shield structure adjustably coupled to the housing to provide electromagnetic shielding for the imaging region in an amount that can be varied by adjusting the at least one electromagnetic shield structure about the imaging region.

BRIEF DESCRIPTION OF THE DRAWINGS

Various aspects and embodiments of the disclosed technology will be described with reference to the following figures. It should be appreciated that the figures are not necessarily drawn to scale.

FIG. 1 illustrates exemplary components of a magnetic resonance imaging system;

FIGS. 2A and 2B illustrate a B₀ magnet comprising a plurality of electromagnets, in accordance with some embodiments;

FIG. 3A illustrates a B₀ magnet comprising a plurality of permanent magnets, in accordance with some embodiments;

FIG. 3B illustrates a top view of an exemplary configuration of permanent magnet rings forming, in part, the B₀ magnet illustrated in FIG. 3A;

FIGS. 4A and 4B illustrate an exemplary ring of permanent magnets for a B₀ magnet, in accordance with some embodiments;

FIGS. 5A-C illustrate exemplary dimensions for permanent magnet blocks for the permanent magnet ring illustrated in FIGS. 4A and 4B, in accordance with some embodiments;

FIGS. 6A-C illustrate exemplary dimensions for permanent magnet blocks for the permanent magnet ring illustrated in FIGS. 4A and 4B, in accordance with some embodiments;

FIGS. 7A and 7B illustrate a portion of an exemplary ring of permanent magnets for a B₀ magnet, in accordance with some embodiments;

FIGS. 8A-C illustrate exemplary dimensions for permanent magnet blocks for the permanent magnet ring illustrated in FIGS. 7A and 7B, in accordance with some embodiments;

FIGS. 9A-C illustrate exemplary dimensions for permanent magnet blocks for the permanent magnet ring illustrated in FIGS. 7A and 7B, in accordance with some embodiments;

FIGS. 10A-C illustrate exemplary dimensions for permanent magnet blocks for the permanent magnet ring illustrated in FIGS. 7A and 7B, in accordance with some embodiments;

FIGS. 11A and 11B illustrate a portion of an exemplary ring of permanent magnets for a B₀ magnet, in accordance with some embodiments;

FIGS. 12A-C illustrate exemplary dimensions for permanent magnet blocks for the permanent magnet ring illustrated in FIGS. 11A and 11B, in accordance with some embodiments;

FIGS. 13A-C illustrate exemplary dimensions for permanent magnet blocks for the permanent magnet ring illustrated in FIGS. 11A and 11B, in accordance with some embodiments;

FIGS. 14A-C illustrate exemplary dimensions for permanent magnet blocks for the permanent magnet ring illustrated in FIGS. 11A and 11B, in accordance with some embodiments;

FIGS. 15A and 15B illustrate a portion of an exemplary ring of permanent magnets for a B₀ magnet, in accordance with some embodiments;

FIGS. 16A-C illustrate exemplary dimensions for permanent magnet blocks for the permanent magnet ring illustrated in FIGS. 15A and 15B, in accordance with some embodiments;

FIGS. 17A-C illustrate exemplary dimensions for permanent magnet blocks for the permanent magnet ring illustrated in FIGS. 15A and 15B, in accordance with some embodiments;

FIGS. 18A-C illustrate exemplary dimensions for permanent magnet blocks for the permanent magnet ring illustrated in FIGS. 15A and 15B, in accordance with some embodiments;

FIGS. 19A and 19B illustrate a portion of an exemplary ring of permanent magnets for a B₀ magnet, in accordance with some embodiments;

FIGS. 20A-C illustrate exemplary dimensions for permanent magnet blocks for the permanent magnet ring illustrated in FIGS. 19A and 19B, in accordance with some embodiments;

FIGS. 21A-C illustrate exemplary dimensions for permanent magnet blocks for the permanent magnet ring illustrated in FIGS. 19A and 19B, in accordance with some embodiments;

FIGS. 22A-C illustrate exemplary dimensions for permanent magnet blocks for the permanent magnet ring illustrated in FIGS. 19A and 19B, in accordance with some embodiments;

FIGS. 23A and 23B illustrate a portion of an exemplary ring of permanent magnets for a B₀ magnet, in accordance with some embodiments;

FIGS. 24A-C illustrate exemplary dimensions for permanent magnet blocks for the permanent magnet ring illustrated in FIGS. 23A and 23B, in accordance with some embodiments;

FIGS. 25A-C illustrate exemplary dimensions for permanent magnet blocks for the permanent magnet ring illustrated in FIGS. 23A and 23B, in accordance with some embodiments;

FIGS. 26A-C illustrate exemplary dimensions for permanent magnet blocks for the permanent magnet ring illustrated in FIGS. 23A and 23B, in accordance with some embodiments;

FIGS. 27A and 27B illustrate a portion of an exemplary ring of permanent magnets for a B₀ magnet, in accordance with some embodiments;

FIGS. 28A-C illustrate exemplary dimensions for permanent magnet blocks for the permanent magnet ring illustrated in FIGS. 27A and 27B, in accordance with some embodiments;

FIGS. 29A-C illustrate exemplary dimensions for permanent magnet blocks for the permanent magnet ring illustrated in FIGS. 27A and 27B, in accordance with some embodiments;

FIGS. 30A-C illustrate exemplary dimensions for permanent magnet blocks for the permanent magnet ring illustrated in FIGS. 27A and 27B, in accordance with some embodiments;

FIGS. 31A-C illustrate views of an exemplary permanent magnet disk, in accordance with some embodiments;

FIG. 32 illustrates a B0 magnet comprising a plurality of permanent magnets, in accordance with some embodiments;

FIG. 33 illustrates a top view of an exemplary configuration of permanent magnet rings forming, in part, the B₀ magnet illustrated in FIG. 32;

FIGS. 34A and 34B illustrate an exemplary ring of permanent magnet segments for a B₀ magnet, in accordance with some embodiments;

FIGS. 35A and 35B illustrate different views of permanent magnet segments that can be used to form the permanent magnet ring illustrated in FIG. 35C, in accordance with some embodiments;

FIG. 35C illustrates a permanent magnet ring for a B₀ magnet, in accordance with some embodiments;

FIGS. 36A and 36B illustrate different views of permanent magnet segments that can be used to form the permanent magnet ring illustrated in FIG. 36C, in accordance with some embodiments;

FIG. 36C illustrates a permanent magnet ring for a B₀ magnet, in accordance with some embodiments;

FIG. 37A illustrate an exemplary portable MRI system, in accordance with some embodiments;

FIGS. 37B-E illustrate a housing configured to house the power components, power supplies, console and power distribution unit within the footprint of the portable MRI system illustrated in FIG. 37A;

FIGS. 38A and 38B illustrate views of a portable MRI system, in accordance with some embodiments;

FIG. 39A illustrates a portable MRI system performing a scan of the head, in accordance with some embodiments;

FIG. 39B illustrates a portable MRI system performing a scan of the knee, in accordance with some embodiments;

FIG. 40 illustrates another example of a portable MRI system, in accordance with some embodiments;

FIG. 41A-D illustrate exemplary respective examples of a noise reduction system, in accordance with some embodiments;

FIG. 42 is a flowchart of an illustrative process for performing noise reduction, in accordance with some embodiments;

FIG. 43A-C illustrate respective examples of decoupling circuits configured to reduce inductive coupling between radio frequency coils in a multi-coil transmit/receive system, in accordance with some embodiments;

FIGS. 44A-C illustrate a portable MRI system having different amounts of device-level shielding about the imaging region, in accordance with some embodiments;

FIG. 44D illustrates a portable MRI system utilizing a conductive strip to provide electromagnetic shielding for the imaging region, in accordance with some embodiments;

FIGS. 45A-45D illustrate different views of a positioning mechanism, in accordance with some embodiments;

FIGS. 46A and 46B illustrate exemplary components of a positioning mechanism, in accordance with some embodiments; and

FIGS. 47A and 47B illustrate a portable low-field MRI system, in accordance with some embodiments.

DETAILED DESCRIPTION

The MRI scanner market is overwhelmingly dominated by high-field systems, and particularly for medical or clinical MRI applications. As discussed above, the general trend in medical imaging has been to produce MRI scanners with increasingly greater field strengths, with the vast majority of clinical MRI scanners operating at 1.5 T or 3 T, with higher field strengths of 7 T and 9 T used in research settings. As used herein, “high-field” refers generally to MRI systems presently in use in a clinical setting and, more particularly, to MRI systems operating with a main magnetic field (i.e., a B₀ field) at or above 1.5 T, though clinical systems operating between 0.5 T and 1.5 T are often also characterized as “high-field.” Field strengths between approximately 0.2 T and 0.5 T have been characterized as “mid-field” and, as field strengths in the high-field regime have continued to increase, field strengths in the range between 0.5 T and 1T have also been characterized as mid-field. By contrast, “low-field” refers generally to MRI systems operating with a B₀ field of less than or equal to approximately 0.2 T, though systems having a B₀ field of between 0.2 T and approximately 0.3 T have sometimes been characterized as low-field as a consequence of increased field strengths at the high end of the high-field regime. Within the low-field regime, low-field MRI systems operating with a B₀ field of less than 0.1 T are referred to herein as “very low-field” and low-field MRI systems operating with a B₀ field of less than 10 mT are referred to herein as “ultra-low field.”

As discussed above, conventional MRI systems require specialized facilities. An electromagnetically shielded room is required for the MRI system to operate and the floor of the room must be structurally reinforced. Additional rooms must be provided for the high-power electronics and the scan technician's control area. Secure access to the site must also be provided. In addition, a dedicated three-phase electrical connection must be installed to provide the power for the electronics that must, in turn, be cooled by a chilled water supply and additional HVAC capacity must be provided. These site requirements are not only costly, but significantly limit the locations where MRI systems can be deployed. Conventional clinical MRI scanners also require substantial expertise to both operate and maintain. These highly trained technicians and service engineers add large on-going operational costs to operating an MRI system. Conventional MRI, as a result, is cost prohibitive and severely limited in accessibility, preventing MRI from being a widely available diagnostic tool capable of delivering a wide range of clinical imaging solutions wherever and whenever needed. Patient must visit one of a limited number of facilities at a time and place scheduled in advance, preventing MRI from being used in numerous medical applications for which it is uniquely efficacious in assisting with diagnosis, surgery, patient monitoring and the like.

As discussed above, high-field MRI systems require specially adapted facilities to accommodate the size, weight, power consumption and shielding requirements of these systems. For example, a 1.5 T MRI system typically weighs between 4-10 tons and a 3 T MRI system typically weighs between 8-20 tons. In addition, high-field MRI systems generally require significant amounts of heavy and expensive shielding. Many mid-field scanners are even heavier, weighing between 10-20 tons due, in part, to the use of very large permanent magnets and/or yokes. Commercially available low-field MRI systems (e.g., operating with a B₀ magnetic field of 0.2 T) are also typically in the range of 10 tons or more due the large of amounts of ferromagnetic material used to generate the B₀ field, with additional tonnage in shielding. To accommodate this heavy equipment, rooms (which typically have a minimum size of 30-50 square meters) have to be built with reinforced flooring (e.g., concrete flooring), and must be specially shielded to prevent electromagnetic radiation from interfering with operation of the MRI system. Thus, available clinical MRI systems are immobile and require the significant expense of a large, dedicated space within a hospital or facility, and in addition to the considerable costs of preparing the space for operation, require further additional on-going costs in expertise in operating and maintaining the system.

In addition, currently available MRI systems typically consume large amounts of power. For example, common 1.5 T and 3 T MRI systems typically consume between 20-40 kW of power during operation, while available 0.5 T and 0.2 T MRI systems commonly consume between 5-20 kW. When discussing power consumption herein, average power consumed over an interval of interest will be referenced unless otherwise specified. For example, the 20-40 kW referred to above indicates the average power consumed by conventional MRI systems during the course of image acquisition, which may include relatively short periods of peak power consumption that significantly exceeds the average power consumption (e.g., when the gradient coils and/or RF coils are pulsed over relatively short periods of the pulse sequence). Intervals of peak (or large) power consumption are typically addressed via power storage elements (e.g., capacitors) of the MRI system itself. Thus, the average power consumption is the more relevant property as it generally determines the type of power connection needed to operate the device. Accordingly, available clinical MRI systems must have dedicated power sources as well, typically requiring a dedicated three-phase connection to the grid for operating the MRI system. Additional electronics are then needed to convert the three-phase power into single-phase power utilized by the MRI system. The many physical requirements of deploying conventional clinical MRI systems creates a significant problem of availability and severely restricts the clinical applications for which MRI can be utilized.

Accordingly, the many requirements of high-field MRI render installations prohibitive in many situations, limiting their deployment to large institutional hospitals or specialized facilities and generally restricting their use to tightly scheduled appointments, requiring the patient to visit dedicated facilities at times scheduled in advance. Thus, the many restrictions on high field MRI prevent MRI from being fully utilized as an imaging modality. Despite the drawbacks of high-field MRI mentioned above, the appeal of the significant increase in SNR at higher fields continues to drive the industry to higher and higher field strengths for use in clinical and medical MRI applications, further increasing the cost and complexity of MRI scanners, and further limiting their availability and preventing their use as a general-purpose and/or generally-available imaging solution.

The low SNR of MR signals produced in the low-field regime (particularly in the very low-field regime) has prevented the development of a relatively low cost, low power and/or portable MRI system. Conventional “low-field” MRI systems operate at the high end of what is typically characterized as the low-field range (e.g., clinically available low-field systems have a floor of approximately 0.2 T) to achieve useful images. Though somewhat less expensive then high-field MRI systems, conventional low-field MRI systems share many of the same drawbacks. In particular, conventional low-field MRI systems are large, fixed and immobile installments, consume substantial power (requiring dedicated three-phase power hook-ups) and require specially shielded rooms and large dedicated spaces. The challenges of low-field MRI have prevented the development of relatively low cost, low power and/or portable MRI systems that can produce useful images.

The inventors have developed techniques enabling portable, low-field, low power and/or lower-cost MRI systems that can improve the wide-scale deployability of MRI technology in a variety of environments beyond the current MRI installments at hospitals and research facilities. As a result, MRI can be deployed in emergency rooms, small clinics, doctor's offices, in mobile units, in the field, etc. and may be brought to the patient (e.g., bedside) to perform a wide variety of imaging procedures and protocols. Some embodiments include very low-field MRI systems (e.g., 0.1 T, 50 mT, 20 mT, etc.) that facilitate portable, low-cost, low-power MRI, significantly increasing the availability of MRI in a clinical setting.

There are numerous challenges to developing a clinical MRI system in the low-field regime. As used herein, the term clinical MRI system refers to an MRI system that produces clinically useful images, which refers to an images having sufficient resolution and adequate acquisition times to be useful to a physician or clinician for its intended purpose given a particular imaging application. As such, the resolutions/acquisition times of clinically useful images will depend on the purpose for which the images are being obtained. Among the numerous challenges in obtaining clinically useful images in the low-field regime is the relatively low SNR. In particular, the relationship between SNR and B₀ field strength is approximately B₀ ^(5/4) at field strength above 0.2 T and approximately B₀ ^(3/2) at field strengths below 0.1 T. As such, the SNR drops substantially with decreases in field strength with even more significant drops in SNR experienced at very low field strength. This substantial drop in SNR resulting from reducing the field strength is a significant factor that has prevented development of clinical MRI systems in the very low-field regime. In particular, the challenge of the low SNR at very low field strengths has prevented the development of a clinical MRI system operating in the very low-field regime. As a result, clinical MRI systems have been limited to approximately the 0.2 T range and above, which systems are large, heavy (and costly) MRI systems that consume significant power, generally requiring fixed dedicated spaces (or shielded tents) and dedicated power sources.

The inventors have developed low-field and very low-field MRI systems capable of producing clinically useful images, allowing for the development of portable, low cost and easy to use MRI systems not achievable using state of the art technology. According to some embodiments, an MRI system can be transported to the patient to provide a wide variety of diagnostic, surgical, monitoring and/or therapeutic procedures, generally, whenever and wherever needed.

FIG. 1 is a block diagram of typical components of a MRI system 100. In the illustrative example of FIG. 1, MRI system 100 comprises computing device 104, controller 106, pulse sequences store 108, power management system 110, and magnetics components 120. It should be appreciated that system 100 is illustrative and that a MRI system may have one or more other components of any suitable type in addition to or instead of the components illustrated in FIG. 1. However, a MRI system will generally include these high level components, though the implementation of these components for a particular MRI system may differ vastly, as discussed in further detail below.

As illustrated in FIG. 1, magnetics components 120 comprise B₀ magnet 122, shim coils 124, RF transmit and receive coils 126, and gradient coils 128. Magnet 122 may be used to generate the main magnetic field B₀. Magnet 122 may be any suitable type or combination of magnetics components that can generate a desired main magnetic B₀ field. As discussed above, in the high field regime, the B₀ magnet is typically formed using superconducting material generally provided in a solenoid geometry, requiring cryogenic cooling systems to keep the B₀ magnet in a superconducting state. Thus, high-field B₀ magnets are expensive, complicated and consume large amounts of power (e.g., cryogenic cooling systems require significant power to maintain the extremely low temperatures needed to keep the B₀ magnet in a superconducting state), require large dedicated spaces, and specialized, dedicated power connections (e.g., a dedicated three-phase power connection to the power grid). Conventional low-field B₀ magnets (e.g., B₀ magnets operating at 0.2 T) are also often implemented using superconducting material and having these same general requirements. Other conventional low-field B₀ magnets are implemented using permanent magnets, which in order to produce the field strengths to which conventional low-field systems are limited to (e.g., between 0.2 T and 0.3 T) due to the inability to acquire useful images at lower field strengths, need to be very large magnets weighing 5-20 tons. Thus, the B₀ magnet of conventional MRI systems alone prevents both portability and affordability.

Gradient coils 128 may be arranged to provide gradient fields and, for example, may be arranged to generate gradients in the B₀ field in three substantially orthogonal directions (X, Y, Z). Gradient coils 128 may be configured to encode emitted MR signals by systematically varying the B₀ field (the B₀ field generated by magnet 122 and/or shim coils 124) to encode the spatial location of received MR signals as a function of frequency or phase. For example, gradient coils 128 may be configured to vary frequency or phase as a linear function of spatial location along a particular direction, although more complex spatial encoding profiles may also be provided by using nonlinear gradient coils. For example, a first gradient coil may be configured to selectively vary the B₀ field in a first (X) direction to perform frequency encoding in that direction, a second gradient coil may be configured to selectively vary the B₀ field in a second (Y) direction substantially orthogonal to the first direction to perform phase encoding, and a third gradient coil may be configured to selectively vary the B₀ field in a third (Z) direction substantially orthogonal to the first and second directions to enable slice selection for volumetric imaging applications. As discussed above, conventional gradient coils also consume significant power, typically operated by large, expensive gradient power sources, as discussed in further detail below.

MRI is performed by exciting and detecting emitted MR signals using transmit and receive coils, respectively (often referred to as radio frequency (RF) coils). Transmit/receive coils may include separate coils for transmitting and receiving, multiple coils for transmitting and/or receiving, or the same coils for transmitting and receiving. Thus, a transmit/receive component may include one or more coils for transmitting, one or more coils for receiving and/or one or more coils for transmitting and receiving. Transmit/receive coils are also often referred to as Tx/Rx or Tx/Rx coils to generically refer to the various configurations for the transmit and receive magnetics component of an MRI system. These terms are used interchangeably herein. In FIG. 1, RF transmit and receive coils 126 comprise one or more transmit coils that may be used to generate RF pulses to induce an oscillating magnetic field B₁. The transmit coil(s) may be configured to generate any suitable types of RF pulses.

Power management system 110 includes electronics to provide operating power to one or more components of the low-field MRI system 100. For example, as discussed in more detail below, power management system 110 may include one or more power supplies, gradient power components, transmit coil components, and/or any other suitable power electronics needed to provide suitable operating power to energize and operate components of MRI system 100. As illustrated in FIG. 1, power management system 110 comprises power supply 112, power component(s) 114, transmit/receive switch 116, and thermal management components 118 (e.g., cryogenic cooling equipment for superconducting magnets). Power supply 112 includes electronics to provide operating power to magnetic components 120 of the MRI system 100. For example, power supply 112 may include electronics to provide operating power to one or more B₀ coils (e.g., B₀ magnet 122) to produce the main magnetic field for the low-field MRI system. Transmit/receive switch 116 may be used to select whether RF transmit coils or RF receive coils are being operated.

Power component(s) 114 may include one or more RF receive (Rx) pre-amplifiers that amplify MR signals detected by one or more RF receive coils (e.g., coils 126), one or more RF transmit (Tx) power components configured to provide power to one or more RF transmit coils (e.g., coils 126), one or more gradient power components configured to provide power to one or more gradient coils (e.g., gradient coils 128), and one or more shim power components configured to provide power to one or more shim coils (e.g., shim coils 124).

In conventional MRI systems, the power components are large, expensive and consume significant power. Typically, the power electronics occupy a room separate from the MRI scanner itself. The power electronics not only require substantial space, but are expensive complex devices, that consume substantial power and require wall mounted racks to be supported. Thus, the power electronics of conventional MRI systems also prevent portability and affordable of MRI.

As illustrated in FIG. 1, MRI system 100 includes controller 106 (also referred to as a console) having control electronics to send instructions to and receive information from power management system 110. Controller 106 may be configured to implement one or more pulse sequences, which are used to determine the instructions sent to power management system 110 to operate the magnetic components 120 in a desired sequence (e.g., parameters for operating the RF transmit and receive coils 126, parameters for operating gradient coils 128, etc.). As illustrated in FIG. 1, controller 106 also interacts with computing device 104 programmed to process received MR data. For example, computing device 104 may process received MR data to generate one or more MR images using any suitable image reconstruction process(es). Controller 106 may provide information about one or more pulse sequences to computing device 104 for the processing of data by the computing device. For example, controller 106 may provide information about one or more pulse sequences to computing device 104 and the computing device may perform an image reconstruction process based, at least in part, on the provided information. In conventional MRI systems, computing device 104 typically includes one or more high performance work-stations configured to perform computationally expensive processing on MR data relatively rapidly. Such computing devices are relatively expensive equipment on their own.

As should be appreciated from the foregoing, currently available clinical MRI systems (including high-field, mid-field and low-field systems) are large, expensive, fixed installations requiring substantial dedicated and specially designed spaces, as well as dedicated power connections. The inventors have developed low-field, including very-low field, MRI systems that are lower cost, lower power and/or portable, significantly increasing the availability and applicability of MRI. According to some embodiments, a portable MRI system is provided, allowing an MRI system to be brought to the patient and utilized at locations where it is needed.

As discussed above, some embodiments include an MRI system that is portable, allowing the MRI device to be moved to locations in which it is needed (e.g., emergency and operating rooms, primary care offices, neonatal intensive care units, specialty departments, emergency and mobile transport vehicles and in the field). There are numerous challenges that face the development of a portable MRI system, including size, weight, power consumption and the ability to operate in relatively uncontrolled electromagnetic noise environments (e.g., outside a specially shielded room). As discussed above, currently available clinical MRI systems range from approximately 4-20 tons. Thus, currently available clinical MRI systems are not portable because of the sheer size and weight of the imaging device itself, let alone the fact that currently available systems also require substantial dedicated space, including a specially shielded room to house the MRI scanner and additional rooms to house the power electronics and the technician control area, respectively. The inventors have developed MRI systems of suitable weight and size to allow the MRI system to be transported to a desired location, some examples of which are discussed in further detail below.

A further aspect of portability involves the capability of operating the MRI system in a wide variety of locations and environments. As discussed above, currently available clinical MRI scanners are required to be located in specially shielded rooms to allow for correct operation of the device and is one (among many) of the reasons contributing to the cost, lack of availability and non-portability of currently available clinical MRI scanners. Thus, to operate outside of a specially shielded room and, more particularly, to allow for generally portable, cartable or otherwise transportable MRI, the MRI system must be capable of operation in a variety of noise environments. The inventors have developed noise suppression techniques that allow the MRI system to be operated outside of specially shielded rooms, facilitating both portable/transportable MRI as well as fixed MRI installments that do not require specially shielded rooms. While the noise suppression techniques allow for operation outside specially shielded rooms, these techniques can also be used to perform noise suppression in shielded environments, for example, less expensive, loosely or ad-hoc shielding environments, and can be therefore used in conjunction with an area that has been fitted with limited shielding, as the aspects are not limited in this respect.

A further aspect of portability involves the power consumption of the MRI system. As also discussed above, current clinical MRI systems consume large amounts of power (e.g., ranging from 20 kW to 40 kW average power consumption during operation), thus requiring dedicated power connections (e.g., dedicated three-phase power connections to the grid capable of delivering the required power). The requirement of a dedicated power connection is a further obstacle to operating an MRI system in a variety of locations other than expensive dedicated rooms specially fitted with the appropriate power connections. The inventors have developed low power MRI systems capable of operating using mains electricity such as a standard wall outlet (e.g., 120V/20 A connection in the U.S.) or common large appliance outlets (e.g., 220-240V/30 A), allowing the device to be operated anywhere common power outlets are provided. The ability to “plug into the wall” facilitates both portable/transportable MRI as well as fixed MRI system installations without requiring special, dedicated power such as a three-phase power connection.

As discussed above, a significant contributor to the size, cost and power consumption of conventional MRI systems are the power electronics for powering the magnetics components of the MRI system. The power electronics for conventional MRI systems often require a separate room, are expensive and consume significant power to operate the corresponding magnetics components. In particular, the gradient coils and thermal management systems utilized to cool the gradient coils alone generally require dedicated power connections and prohibit operation from standard wall outlets. The inventors have developed low power, low noise gradient power sources capable of powering the gradient coils of an MRI system that can, in accordance with some embodiments, be housed in the same portable, cartable or otherwise transportable apparatus as the magnetics components of the MRI system. According to some embodiments, the power electronics for powering the gradient coils of an MRI system consume less than 50 W when the system is idle and between 100-200 W when the MRI system is operating (i.e., during image acquisition). The inventors have developed power electronics (e.g., low power, low noise power electronics) to operate a portable low field MRI system that all fit within the footprint of the portable MRI scanner. According to some embodiments, innovative mechanical design has enabled the development of an MRI scanner that is maneuverable within the confines of a variety of clinical environments in which the system is needed.

At the core of developing a low power, low cost and/or portable MRI system is the reduction of the field strength of the B₀ magnet, which can facilitate a reduction in size, weight, expense and power consumption. However, as discussed above, reducing the field strength has a corresponding and significant reduction in SNR. This significant reduction in SNR has prevented clinical MRI systems from reducing the field strength below the current floor of approximately 0.2 T, which systems remains large, heavy, expensive fixed installations requiring specialized and dedicated spaces. While some systems have been developed that operate between 0.1 T and 0.2 T, these systems are often specialized devices for scanning extremities such as the hand, arm or knee. The inventors have developed MRI systems operating in the low-field and very-low field capable of acquiring clinically useful images. Some embodiments include highly efficient pulse sequences in producing MR signals and/or some embodiments include optimized receive coils for detecting emitted MR signals, examples of which are discussed in further detail below. Noise suppression, cancellation and/or avoidance techniques facilitate extraction of as much signal as possible and also allow operation of the MRI system in unshielded or partially shielded environments, as discussed above and in further detail below.

Another aspect of increasing the availability of MRI is to make MRI affordable. The development of a portable low-field MRI system by the inventors eliminates many of the costs associated with conventional clinical MRI systems, including expensive superconducting materials and cryogenic cooling systems, expensive site preparation of large and complex dedicated facilities, highly trained personnel to operate and maintain the system to name a few. In addition, the inventors have developed further cost reduction techniques and designs, including, according to some embodiments, integrated power electronics, designs that reduce materials, optimize or otherwise minimize the use of expensive materials and/or reduce production costs. The inventors have developed automated shimming techniques to allow for correction of field inhomogeneity of the B₀ magnet after manufacture, reducing the cost of both production and post-production processes.

According to some embodiments, designs developed by the inventors also reduce the cost and complexity of operating and maintaining the MRI scanner. For example, conventional clinical MRI systems require significant expertise to both operate and maintain, resulting in significant on-going costs of these systems. The inventors have developed easy-to-use an MRI systems that allow minimally trained or untrained personnel to operate and/or maintain the system. According to some embodiments, automated setup processes allow the MRI scanner to automatically probe and adapt to its environment to prepare for operation. Network connectivity allows the MRI system to be operated from a mobile device such as a tablet, notepad or smart phone with easy-to-use interfaces configured to automatically run desired scanning protocols. Acquired images are immediately transferred to a secure cloud server for data sharing, telemedicine and/or deep learning.

Following below are more detailed descriptions of various concepts related to, and embodiments of, lower cost, lower power and/or portable low-field MRI. It should be appreciated that the embodiments described herein may be implemented in any of numerous ways. Examples of specific implementations are provided below for illustrative purposes only. It should be appreciated that the embodiments and the features/capabilities provided may be used individually, all together, or in any combination of two or more, as aspects of the technology described herein are not limited in this respect.

A significant contributor to the high cost, size, weight and power consumption of high-field MRI is the B₀ magnet itself along with the apparatus required to power the B₀ magnet and to perform thermal management thereof. In particular, to produce the field strengths characteristic of high-field MRI, the B₀ magnet is typically implemented as an electromagnet configured in a solenoid geometry using superconducting wires that need a cryogenic cooling system to keep the wires in a superconducting state. Not only is the superconducting material itself expensive, but the cryogenic equipment to maintain the superconducting state is also expensive and complex.

The inventors have recognized that the low-field context allows for B₀ magnet designs not feasible in the high-field regime. For example, due at least in part to the lower field strengths, superconducting material and the corresponding cryogenic cooling systems can be eliminated. Due in part to the low-field strengths, B₀ electromagnets constructed using non-superconducting material (e.g., copper) may be employed in the low-field regime. However, such electromagnets still may consume relatively large amounts of power during operation. For example, operating an electromagnet using a copper conductor to generate a magnetic field of 0.2 T or more requires a dedicated or specialized power connection (e.g., a dedicated three-phase power connection). The inventors have developed MRI systems that can be operated using mains electricity (i.e., standard wall power), allowing the MRI system to be powered at any location having common power connection, such as a standard wall outlet (e.g., 120V/20 A connection in the U.S.) or common large appliance outlets (e.g., 220-240V/30 A). Thus, a low-power MRI system facilitates portability and availability, allowing an MRI system to be operated at locations where it is needed (e.g., the MRI system can be brought to the patient instead of vice versa), examples of which are discussed in further detail below. In addition, operating from standard wall power eliminates the electronics conventionally needed to convert three-phase power to single-phase power and to smooth out the power provided directly from the grid. Instead, wall power can be directly converted to DC and distributed to power the components of the MRI system.

FIGS. 2A and 2B illustrate a B₀ magnet formed using an electromagnet and a ferromagnetic yoke. In particular, B₀ magnet 200 is formed in part by an electromagnet 210 arranged in a bi-planar geometry comprising electromagnetic coils 212 a and 212 b on an upper side and electromagnetic coils 214 a and 214 b on a lower side of B₀ magnet 200. According to some embodiments, the coils forming electromagnet 210 may be formed from a number of turns of a copper wire or copper ribbon, or any other conductive material suitable for producing a magnetic field when operated (e.g., when electrical current is driven through the conductor windings). While the exemplary electromagnet illustrated in FIGS. 2A and 2B comprises two pairs of coils, an electromagnet may be formed using any number of coils in any configuration, as the aspects are not limited in this respect. The electromagnetic coils forming electromagnet 210 may be formed, for example, by winding a conductor 213 (e.g., a copper ribbon, wire, paint, etc.) about a fiberglass ring 217. For example, conductor 213 may be a suitable insulated copper wire, or alternatively, conductor 213 may be a copper ribbon wound in conjunction with an insulating layer (e.g., a Mylar layer) to electrically isolate the multiple windings of the coil. A connector 219 may be provided to allow for a power connection to provide current to operate coils 214 a and 214 b in series. A similar connector on the upper side of the electromagnet (not visible in FIGS. 2A and 2B) may be provided to operate coils 212 a and 212 b.

It should be appreciated that the electromagnetic coils may be formed from any suitable material and dimensioned in any suitable way so as to produce or contribute to a desired B₀ magnetic field, as the aspects are not limited for use with any particular type of electromagnet. As one non-limiting example that may be suitable to form, in part, an electromagnet (e.g., electromagnet 210), an electromagnetic coil may be constructed using copper ribbon and mylar insulator having 155 turns to form an inner diameter of approximately 23-27 inches (e.g., approximately 25 inches), an outer diameter of approximately 30-35 inches (e.g., 32 inches). However, different material and/or different dimensions may be used to construct an electromagnetic coil having desired characteristics, as the aspects are not limited in this respect. The upper and lower coil(s) may be positioned to provide a distance of approximately 10-15 inches (e.g., approximately 12.5 inches) between the lower coil on the upper side and the upper coil on the lower side. It should be appreciated that the dimensions will differ depending on the desired characteristics including, for example, field strength, field of view, etc.

In the exemplary B₀ magnet illustrated in FIGS. 2A and 2B, each coil pair 212 and 214 is separated by thermal management components 230 a and 230 b, respectively, to transfer heat produced by the electromagnetic coils and gradient coils (not illustrated in FIGS. 2A and 2B) away from the magnets to provide thermal management for the MRI device. In particular, thermal management components 230 a and 230 b may comprise a cooling plate having conduits that allow coolant to be circulated through the cooling plate to transfer heat away from the magnets. The cooling plate 230 a, 230 b may be constructed to reduce or eliminate eddy currents induced by operating the gradient coils that can produce electromagnetic fields that disrupt the B₀ magnetic field produced by the B0 magnet 200. For example, thermal management components 230 a and 230 b may be the same or similar to any of the thermal management components described in U.S. application Ser. No. 14/846,042 entitled “Thermal Management Methods and Apparatus,” filed on Sep. 4, 2015, which is incorporated by reference herein in its entirety. According to some embodiments, thermal management components may be eliminated, as discussed in further detail below.

B0 magnet 200 further comprises a yoke 220 that is magnetically coupled to the electromagnet to capture magnetic flux that, in the absence of yoke 220, would be lost and not contribute to the flux density in the region of interest between the upper and lower electromagnetic coils. In particular, yoke 220 forms a “magnetic circuit” connecting the coils on the upper and lower side of the electromagnet so as to increase the flux density in the region between the coils, thus increasing the field strength within the imaging region (also referred to as the field of view) of the B₀ magnet. The imaging region or field of view defines the volume in which the B₀ magnetic field produced by a given B0 magnet is suitable for imaging. More particularly, the imaging region or field of view corresponds to the region for which the B₀ magnetic field is sufficiently homogeneous at a desired field strength that detectable MR signals are emitted by an object positioned therein in response to application of radio frequency excitation (e.g., a suitable radio frequency pulse sequence). Yoke 220 comprises frame 222 and plates 224 a, 224 b, which may be formed using any suitable ferromagnetic material (e.g., iron, steel, etc.). Plates 224 a, 224 b collect magnetic flux generated by the coil pairs of electromagnet 210 and directs it to frame 222 which, in turn, returns the flux back to the opposing coil pair, thereby increasing, by up to a factor of two, the magnetic flux density in the imaging region between the coil pairs (e.g., coil pair 212 a, 212 b and coil pair 214 a, 214 b) for the same amount of operating current provided to the coils. Thus, yoke 220 can be used to produce a higher B₀ field (resulting in higher SNR) without a corresponding increase in power requirements, or yoke 220 can be used to lower the power requirements of B₀ magnet 200 for a given B₀ field.

According to some embodiments, the material used for portions of yoke 220 (i.e., frame 222 and/or plates 224 a, 224 b) is steel, for example, a low-carbon steel, silicon steel, cobalt steel, etc. According to some embodiments, gradient coils (not shown in FIGS. 2A, 2B) of the MRI system are arranged in relatively close proximity to plates 224 a, 224 b inducing eddy currents in the plates. To mitigate, plates 224 a, 224 b and/or frame 222 may be constructed of silicon steel, which is generally more resistant to eddy current production than, for example, low-carbon steel. It should be appreciated that yoke 220 may be constructed using any ferromagnetic material with sufficient magnetic permeability and the individual parts (e.g., frame 222 and plates 224 a, 224 b) may be constructed of the same or different ferromagnetic material, as the techniques of increasing flux density is not limited for use with any particular type of material or combination of materials. Furthermore, it should be appreciated that yoke 220 can be formed using different geometries and arrangements. In particular, yoke 220 may be, for example, arranged in any of the geometries illustrated in U.S. patent application Ser. No. 14/846,255 ('255 Application), filed Sep. 4, 2015 and entitled “Ferromagnetic Augmentation For Magnetic Resonance Imaging.”

It should be appreciated that the yoke 220 may be made of any suitable material and may be dimensioned to provide desired magnetic flux capture while satisfying other design constraints such as weight, cost, magnetic properties, etc. As an example, the frame of the yoke (e.g., frame 222) may be formed of a low-carbon steel of less than 0.2% carbon or silicon steel, with the long beam(s) having a length of approximately 38 inches, a width of approximately 8 inches, and a thickness (depth) of approximately 2 inches, and the short beam(s) having a length of approximately 19 inches, a width of approximately 8 inches and a thickness (depth of approximately 2 inches. The plates (e.g., plates 224 a and 224 b) may be formed from a low-carbon steel of less than 0.2% carbon or silicon steel and have a diameter of approximately 30-35 inches (e.g., approximately 32 inches). However, the above provided dimensions and materials are merely exemplary of a suitable embodiment of a yoke that can be used to capture magnetic flux generated by an electromagnet.

As an example of the improvement achieved via the use of yoke 220, operating electromagnet 210 to produce a B₀ magnetic field of approximately 20 mT without yoke 220 consumes about 5 kW, while producing the same 20 mT B₀ magnetic field with yoke 220 consumes about 750 W of power. Operating electromagnet 210 with the yoke 220, a B₀ magnetic field of approximately 40 mT may be produced using 2 kW of power and a B₀ magnetic field of approximately 50 mT may be produced using approximately 3 kW of power. Thus, the power requirements can be significantly reduced by use of yoke 220 allowing for operation of a B₀ magnet without a dedicated three-phase power connection. For example, mains electrical power in the United States and most of North America is provided at 120V and 60 Hz and rated at 15 or 20 amps, permitting utilization for devices operating below 1800 and 2400 W, respectively. Many facilities also have 220-240 VAC outlets with 30 amp ratings, permitting devices operating up to 7200 W to be powered from such outlets. According to some embodiments, a low-field MRI system utilizing a B₀ magnet comprising an electromagnet and a yoke (e.g., B₀ magnet 200) is configured to be powered via a standard wall outlet, as discussed in further detail below. According to some embodiments, a low-field MRI system utilizing a B₀ magnet comprising an electromagnet and a yoke (e.g., B₀ magnet 200) is configured to be powered via a 220-240 VAC outlet, as also discussed in further detail below.

Referring again to FIGS. 2A and 2B, exemplary B₀ magnet 210 further comprises shim rings 240 a, 240 b and shim disks 242 a, 242 b configured to augment the generated B₀ magnetic field to improve homogeneity in the field of view (e.g., in the region between the upper and lower coils of the electromagnet where the B₀ field is suitable for sufficient MR signal production), as best seen in FIG. 2B in which the lower coils have been removed. In particular, shim rings 240 and shim disk 242 are dimensioned and arranged to increase the uniformity of the magnetic field generated by the electromagnet at least within the field of view of the B₀ magnet. In particular, the height, thickness and material of shim rings 240 a, 240 b and the diameter, thickness and material of shim disks 242 a, 242 b may be chosen so as to achieve a B₀ field of suitable homogeneity. For example, the shim disk may be provided with a diameter of approximately 5-6 inches and a width of approximately 0.3-0.4 inches. A shim ring may be formed from a plurality of circular arc segments (e.g., 8 circular arc segments) each having a height of approximately 20-22 inches, and a width of approximately 2 inches to form a ring having an inner diameter of approximately between 21-22 inches and approximately between 23-24 inches.

The weight of the B₀ magnet is a significant portion of the overall weight of the MRI system which, in turn, impacts the portability of the MRI system. In embodiments that primarily use low carbon and/or silicon steel for the yoke and shimming components, an exemplary B₀ magnet 200 dimensioned similar to that described in the foregoing may weigh approximately 550 kilograms. According to some embodiments, cobalt steel (CoFe) may be used as the primary material for the yoke (and possibly the shim components), potentially reducing the weight of B₀ magnet 200 to approximately 450 Kilograms. However, CoFe is generally more expensive than, for example, low carbon steel, driving up the cost of the system. Accordingly, in some embodiments, select components may be formed using CoFe to balance the tradeoff between cost and weight arising from its use. Using such exemplary B₀ magnets a portable, cartable or otherwise transportable MRI system may be constructed, for example, by integrating the B₀ magnet within a housing, frame or other body to which castors, wheels or other means of locomotion can be attached to allow the MRI system to be transported to desired locations (e.g., by manually pushing the MRI system and/or including motorized assistance). As a result, an MRI system can be brought to the location in which it is needed, increasing its availability and use as a clinical instrument and making available MRI applications that were previously not possible.

The primary contributor to the overall power consumption of a low-field MRI system employing a B₀ magnet such as B₀ magnet 200 is the electromagnet (e.g., electromagnet 210). For example, in some embodiments, the electromagnet may consume 80% or more of the power of the overall MRI system. To significantly reduce the power requirements of the MRI system, the inventors have developed B₀ magnets that utilize permanent magnets to produce and/or contribute to the B₀ electromagnetic field. According to some embodiments, B₀ electromagnets are replaced with permanent magnets as the main source of the B₀ electromagnetic field. A permanent magnet refers to any object or material that maintains its own persistent magnetic field once magnetized. Materials that can be magnetized to produce a permanent magnet are referred to herein as ferromagnetic and include, as non-limiting examples, iron, nickel, cobalt, neodymium (NdFeB) alloys, samarium cobalt (SmCo) alloys, alnico (AlNiCo) alloys, strontium ferrite, barium ferrite, etc. Permanent magnet material (e.g., magnetizable material that has been driven to saturation by a magnetizing field) retains its magnetic field when the driving field is removed. The amount of magnetization retained by a particular material is referred to as the material's remanence. Thus, once magnetized, a permanent magnet generates a magnetic field corresponding to its remanence, eliminating the need for a power source to produce the magnetic field.

FIG. 3A illustrates a permanent B₀ magnet, in accordance with some embodiments. In particular, B₀ magnet 300 is formed by permanent magnets 310 a and 310 b arranged in a bi-planar geometry and a yoke 320 that captures electromagnetic flux produced by the permanent magnets and transfers the flux to the opposing permanent magnet to increase the flux density between permanent magnets 310 a and 310 b. Each of permanent magnets 310 a and 310 b are formed from a plurality of concentric permanent magnets. In particular, as visible in FIG. 3, permanent magnetic 310 b comprises an outer ring of permanent magnets 314 a, a middle ring of permanent magnets 314 b, an inner ring of permanent magnets 314 c, and a permanent magnet disk 314 d at the center. Permanent magnet 310 a may comprise the same set of permanent magnet elements as permanent magnet 310 b.

The permanent magnet material used may be selected depending on the design requirements of the system. For example, according to some embodiments, the permanent magnets (or some portion thereof) may be made of NdFeB, which produces a magnetic field with a relatively high magnetic field per unit volume of material once magnetized. According to some embodiments, SmCo material is used to form the permanent magnets, or some portion thereof. While NdFeB produces higher field strengths (and in general is less expensive than SmCo), SmCo exhibits less thermal drift and thus provides a more stable magnetic field in the face of temperature fluctuations. Other types of permanent magnet material(s) may be used as well, as the aspects are not limited in this respect. In general, the type or types of permanent magnet material utilized will depend, at least in part, on the field strength, temperature stability, weight, cost and/or ease of use requirements of a given B₀ magnet implementation.

The permanent magnet rings are sized and arranged to produce a homogenous field of a desired strength in the central region (field of view) between permanent magnets 310 a and 310 b. In the exemplary embodiment illustrated in FIG. 3A, each permanent magnet ring comprises a plurality segments, each segment formed using a plurality of blocks that are stacked in the radial direction and positioned adjacent to one another about the periphery to form the respective ring. The inventors have appreciated that by varying the width (in the direction tangent to the ring) of each permanent magnet, less waste of useful space may be achieved while using less material. For example, the space between stacks that does not produce useful magnetic fields can be reduced by varying the width of the blocks, for example, as function of the radial position of the block, allowing for a closer fit to reduce wasted space and maximize the amount of magnetic field that can be generated in a given space. The dimensions of the blocks may also be varied in any desired way to facilitate the production of a magnetic field of desired strength and homogeneity, as discussed in further detail below.

B₀ magnet 300 further comprises yoke 320 configured and arranged to capture magnetic flux generated by permanent magnets 310 a and 310 b and direct it to the opposing side of the B₀ magnet to increase the flux density in between permanent magnets 310 a and 310 b, increasing the field strength within the field of view of the B₀ magnet. By capturing magnetic flux and directing it to the region between permanent magnets 310 a and 310 b, less permanent magnet material can be used to achieve a desired field strength, thus reducing the size, weight and cost of the B₀ magnet. Alternatively, for given permanent magnets, the field strength can be increased, thus improving the SNR of the system without having to use increased amounts of permanent magnet material. For exemplary B₀ magnet 300, yoke 320 comprises a frame 322 and plates 324 a and 324 b. In a manner similar to that described above in connection with yoke 220, plates 324 a and 324 b capture magnetic flux generated by permanent magnets 310 a and 310 b and direct it to frame 322 to be circulated via the magnetic return path of the yoke to increase the flux density in the field of view of the B₀ magnet. Yoke 320 may be constructed of any desired ferromagnetic material, for example, low carbon steel, CoFe and/or silicon steel, etc. to provide the desired magnetic properties for the yoke. According to some embodiments, plates 324 a and 324 b (and/or frame 322 or portions thereof) may be constructed of silicon steel or the like in areas where the gradient coils could most prevalently induce eddy currents.

Exemplary frame 322 comprises arms 323 a and 323 b that attach to plates 324 a and 324 b, respectively, and supports 325 a and 325 b providing the magnetic return path for the flux generated by the permanent magnets. The arms are generally designed to reduce the amount of material needed to support the permanent magnets while providing sufficient cross-section for the return path for the magnetic flux generated by the permanent magnets. Arms 323 a has two supports within a magnetic return path for the B₀ field produced by the B₀ magnet. Supports 325 a and 325 b are produced with a gap 327 formed between, providing a measure of stability to the frame and/or lightness to the structure while providing sufficient cross-section for the magnetic flux generated by the permanent magnets. For example, the cross-section needed for the return path of the magnetic flux can be divided between the two support structures, thus providing a sufficient return path while increasing the structural integrity of the frame. It should be appreciated that additional supports may be added to the structure, as the technique is not limited for use with only two supports and any particular number of multiple support structures.

As discussed above, exemplary permanent magnets 310 a and 310 b comprise a plurality of rings of permanent magnetic material concentrically arranged with a permanent magnet disk at the center. Each ring may comprise a plurality of stacks of ferromagnetic material to form the respective ring, and each stack may include one or more blocks, which may have any number (including a single block in some embodiments and/or in some of the rings). The blocks forming each ring may be dimensioned and arranged to produce a desired magnetic field. The inventors have recognized that the blocks may be dimensioned in a number of ways to decrease cost, reduce weight and/or improve the homogeneity of the magnetic field produced, as discussed in further detail in connection with the exemplary rings that together form permanent magnets of a B₀ magnet, in accordance with some embodiments.

FIG. 3B illustrates a top-down view of a permanent magnet 310, which may, for example, be used as the design for permanent magnets 310 a and 310 b of B₀ magnet 300 illustrated in FIG. 3A. Permanent magnet 310 comprises concentric rings 310 a, 310 b, and 310 c, each constructed of a plurality of stacks of ferromagnetic blocks, and a ferromagnetic disk 310 d at the center. The direction of the frame of the yoke to which permanent magnet is attached is indicated by arrow 22. In embodiments in which the yoke is not symmetric (e.g., yoke 320), the yoke will cause the magnetic field produced by the permanent magnets for which it captures and focuses magnetic flux to be asymmetric as well, negatively impacting the uniformity of the B₀ magnetic field.

According to some embodiments, the block dimensions are varied to compensate for the effects of the yoke on the magnetic field produced by the permanent magnet. For example, dimensions of blocks in the four regions 315 a, 315 b, 315 c and 315 d labeled in FIG. 3B may be varied depending on which region the respective block is located. In particular, the height of the blocks (e.g., the dimension of the block normal to the plane of the circular magnet 310) may be greater in region 315 c farthest away from the frame than corresponding blocks in region 315 a closest to the frame. Block height can be varied in one or more rings or portions thereof, as the technique of compensating for the effects of the yoke are not limited to varying any particular block, set of blocks and/or any particular dimension. One example of varying block dimension to compensate for yoke effects are discussed in further detail below.

FIGS. 4A and 4B illustrate different views of an inner ring 410 (e.g., ring 310 c illustrated in FIG. 3B), in accordance with some embodiments. Exemplary ring 410 includes a plurality (twelve in FIGS. 4A and 4B) of stacks of two blocks each, thus forming two sub-rings of ferromagnetic blocks (e.g., blocks formed of NdFeB, SmCo, etc.). The inner sub-ring is formed of blocks (e.g., exemplary block 405 b) having a length x, a width y and a height (or depth) z. The outer sub-ring is formed of blocks (e.g., exemplary block 405 a) having a length X, a width y and a height z. As shown, the blocks in the outer sub-ring have a greater length than block in the inner sub-ring (i.e., x<X), reducing the amount of empty space between adjacent block than if the blocks in the outer sub-ring were formed with a length x. As such more of the space in which exemplary ring 410 is contained is occupied by field producing magnetic material, increasing the field strength in the same amount of space. It should be appreciated that the arrangement in exemplary ring 410 is merely illustrative and other arrangements of blocks (e.g., number of stacks and number of blocks within each stack) may be used, as the aspects are not limited in this respect.

FIGS. 5A-C and 6A-C illustrate exemplary dimensions for blocks comprising the inner sub-ring and outer sub-ring of an inner permanent magnet ring (e.g., exemplary dimensions for blocks 405 a and 405 b forming permanent magnet ring 410). In particular, exemplary block 505 illustrated in FIG. 5A (e.g., a block in the inner sub-ring of inner permanent magnet ring 410 or 310 c) may be manufactured to have dimensions as labeled in FIGS. 5B and 5C, where the dimensions are given in millimeters. Exemplary block 605 illustrated in FIG. 6A (e.g., a block in the outer sub-ring of inner permanent magnet ring 410) may be manufactured to have dimensions as labeled in FIGS. 6B and 6C, where the dimensions are also given in millimeters. It should be appreciated that the dimensions of exemplary blocks 505 and 605 are merely illustrative and the dimensions may be selected as desired and are not limited in this respect. Additionally, blocks may be formed using any one or combination of ferromagnetic material, as the aspects are not limited for use with any particular type of magnetic material.

FIGS. 7A and 7B illustrate different views of a portion 715 of middle ring 710 in a quadrant away from the yoke frame (e.g., the portion of ring 310 b in quadrant 315 c illustrated in FIG. 3B), in accordance with some embodiments. Exemplary portion 715 of ring 710 includes a plurality (five in FIGS. 7A and 7B) of stacks of three blocks each, thus forming three sub-rings of ferromagnetic blocks (e.g., blocks formed of NdFeB, SmCo, etc.). The inner sub-ring is formed of blocks (e.g., exemplary block 705 c) having a length x, a width y and a height (or depth) z. The middle sub-ring is formed of blocks (e.g., exemplary block 705 b) having a length X, a width y and a height z. The outer sub-ring is formed of blocks (e.g., exemplary block 705 a) having a length X′, a width y and a height z. As shown, the blocks in the outer sub-ring have a greater length than blocks in the middle sub-ring which, in turn, have a length greater than the blocks in the inner sub-ring (i.e., x<X<X′), reducing the amount of empty space between adjacent block than if the blocks in all sub-rings were formed with a length x. As such more of the space in which exemplary ring 710 is contained is occupied by field producing magnetic material, increasing the field strength in the same amount of space. It should be appreciated that the arrangement in exemplary ring 710 is merely illustrative and other arrangements of blocks (e.g., number of stacks and number of blocks within each stack) may be used, as the aspects are not limited in this respect.

FIGS. 8A-C, 9A-C and 10A-C illustrate exemplary dimensions for blocks comprising the inner, middle and outer sub-rings of a middle permanent magnet ring (e.g., exemplary dimensions for blocks 705 a-c forming permanent magnet ring 710). In particular, exemplary block 805 illustrated in FIG. 8A (e.g., a block in the inner sub-ring of middle permanent magnet ring 710 or 310 b) may be manufactured to have dimensions as labeled in FIGS. 8B and 8C, where the dimensions are given in millimeters. Exemplary block 905 illustrated in FIG. 9A (e.g., a block in the middle sub-ring of middle permanent magnet ring 710 or 310 b) may be manufactured to have dimensions as labeled in FIGS. 9B and 9C, where the dimensions are also given in millimeters. Similarly, exemplary block 1005 illustrated in FIG. 9A (e.g., a block in the outer sub-ring of middle permanent magnet ring 710 or 310 b) may be manufactured to have dimensions as labeled in FIGS. 10B and 10C, where the dimensions are also given in millimeters. It should be appreciated that the dimensions of exemplary blocks 805, 905 and 1005 are merely illustrative and the dimensions may be selected as desired and are not limited in this respect. Additionally, blocks may be formed using any one or combination of ferromagnetic material, as the aspects are not limited for use with any particular type of magnetic material.

FIGS. 11A and 11B illustrate different views of a portion 1115 of middle ring 1110 in quadrant(s) in the middle with respect to the yoke frame (e.g., the portion of ring 310 b in quadrant 315 b and/or 315 d illustrated in FIG. 3B), in accordance with some embodiments. That is, portion 1115 may be used for both middle quadrants, for example, in embodiments where the middle quadrants are equidistant from the yoke frame. Exemplary portion 1115 of ring 1110 includes a plurality (five in FIGS. 11A and 11B) of stacks of three blocks each, thus forming three sub-rings of ferromagnetic blocks (e.g., blocks formed of NdFeB, SmCo, etc.). The inner sub-ring is formed of blocks (e.g., exemplary block 1105 c) having a length x, a width y and a height (or depth) z. The middle sub-ring is formed of blocks (e.g., exemplary block 1105 b) having a length X, a width y and a height z. The outer sub-ring is formed of blocks (e.g., exemplary block 1105 a) having a length X′, a width y and a height z. As shown, the blocks in the outer sub-ring have a greater length than blocks in the middle sub-ring which, in turn, have a length greater than the blocks in the inner sub-ring (i.e., x<X<X′), reducing the amount of empty space between adjacent blocks than if the blocks in all sub-rings were formed with a length x. As such more of the space in which exemplary ring 1110 is contained is occupied by field producing magnetic material, increasing the field strength in the same amount of space. It should be appreciated that the arrangement in exemplary ring 1110 is merely illustrative and other arrangements of blocks (e.g., number of stacks and number of blocks within each stack) may be used, as the aspects are not limited in this respect.

FIGS. 12A-C, 13A-C and 14A-C illustrate exemplary dimensions for blocks comprising the inner, middle and outer sub-rings of a middle permanent magnet ring (e.g., exemplary dimensions for blocks 705 a-c forming permanent magnet ring 710). In particular, exemplary block 1205 illustrated in FIG. 12A (e.g., a block in the inner sub-ring of middle permanent magnet ring 1110 or 310 b) may be manufactured to have dimensions as labeled in FIGS. 12B and 12C, where the dimensions are given in millimeters. Exemplary block 1305 illustrated in FIG. 13A (e.g., a block in the middle sub-ring of middle permanent magnet ring 1110 or 310 b) may be manufactured to have dimensions as labeled in FIGS. 13B and 13C, where the dimensions are also given in millimeters. Similarly, exemplary block 1405 illustrated in FIG. 14A (e.g., a block in the outer sub-ring of middle permanent magnet ring 1110 or 310 b) may be manufactured to have dimensions as labeled in FIGS. 14B and 14C, where the dimensions are also given in millimeters. It should be appreciated that the dimensions of exemplary blocks 1205, 1305 and 1405 are merely illustrative and the dimensions may be selected as desired and are not limited in this respect. Additionally, blocks may be formed using any one or combination of ferromagnetic material, as the aspects are not limited for use with any particular type of magnetic material.

FIGS. 15A and 15B illustrate different views of a portion 1515 of middle ring 1510 in a quadrant nearest the yoke frame (e.g., the portion of ring 310 b in quadrant 315 a illustrated in FIG. 3B), in accordance with some embodiments. Exemplary portion 1515 of ring 1510 includes a plurality (five in FIGS. 15A and 15B) of stacks of three blocks each, thus forming three sub-rings of ferromagnetic blocks (e.g., blocks formed of NdFeB, SmCo, etc.). The inner sub-ring is formed of blocks (e.g., exemplary block 1505 c) having a length x, a width y and a height (or depth) z. The middle sub-ring is formed of blocks (e.g., exemplary block 1505 b) having a length X, a width y and a height z. The outer sub-ring is formed of blocks (e.g., exemplary block 1505 a) having a length X′, a width y and a height z. As shown, the blocks in the outer sub-ring have a greater length than blocks in the middle sub-ring which, in turn, have a length greater than the blocks in the inner sub-ring (i.e., x<X<X′), reducing the amount of empty space between adjacent block than if the blocks in all sub-rings were formed with a length x. As such more of the space in which exemplary ring 1510 is contained is occupied by field producing magnetic material, increasing the field strength in the same amount of space. It should be appreciated that the arrangement in exemplary ring 1510 is merely illustrative and other arrangements of blocks (e.g., number of stacks and number of blocks within each stack) may be used, as the aspects are not limited in this respect.

FIGS. 16A-C, 17A-C and 18A-C illustrate exemplary dimensions for blocks comprising the inner, middle and outer sub-rings of a middle permanent magnet ring (e.g., exemplary dimensions for blocks 1505 a-c forming permanent magnet ring 1510). In particular, exemplary block 1605 illustrated in FIG. 16A (e.g., a block in the inner sub-ring of middle permanent magnet ring 1510 or 310 b) may be manufactured to have dimensions as labeled in FIGS. 16B and 16C, where the dimensions are given in millimeters. Exemplary block 1705 illustrated in FIG. 17A (e.g., a block in the middle sub-ring of middle permanent magnet ring 1510 or 310 b) may be manufactured to have dimensions as labeled in FIGS. 17B and 17C, where the dimensions are also given in millimeters. Similarly, exemplary block 1805 illustrated in FIG. 18A (e.g., a block in the outer sub-ring of middle permanent magnet ring 1510 or 310 b) may be manufactured to have dimensions as labeled in FIGS. 18B and 18C, where the dimensions are also given in millimeters. It should be appreciated that the dimensions of exemplary blocks 1605, 1705 and 1805 are merely illustrative and the dimensions may be selected as desired and are not limited in this respect. Additionally, blocks may be formed using any one or combination of ferromagnetic material, as the aspects are not limited for use with any particular type of magnetic material.

FIGS. 19A and 19B illustrate different views of a portion 1915 of outer ring 1910 in a quadrant away from the yoke frame (e.g., the portion of ring 310 a in quadrant 315 c illustrated in FIG. 3B), in accordance with some embodiments. Exemplary portion 1915 of ring 1910 includes a plurality (nine in FIGS. 19A and 19B) of stacks of three blocks each, thus forming three sub-rings of ferromagnetic blocks (e.g., blocks formed of NdFeB, SmCo, etc.). The inner sub-ring is formed of blocks (e.g., exemplary block 1905 c) having a length x, a width y and a height (or depth) z. The middle sub-ring is formed of blocks (e.g., exemplary block 1905 b) having a length X, a width y and a height z. The outer sub-ring is formed of blocks (e.g., exemplary block 1905 a) having a length X′, a width y and a height z. As shown, the blocks in the outer sub-ring have a greater length than blocks in the middle sub-ring which, in turn, have a length greater than the blocks in the inner sub-ring (i.e., x<X<X′), reducing the amount of empty space between adjacent block than if the blocks in all sub-rings were formed with a length x. As such more of the space in which exemplary ring 1910 is contained is occupied by field producing magnetic material, increasing the field strength in the same amount of space. It should be appreciated that the arrangement in exemplary ring 1910 is merely illustrative and other arrangements of blocks (e.g., number of stacks and number of blocks within each stack) may be used, as the aspects are not limited in this respect.

FIGS. 20A-C, 21A-C and 22A-C illustrate exemplary dimensions for blocks comprising the inner, middle and outer sub-rings of an outer permanent magnet ring (e.g., exemplary dimensions for blocks 1905 a-c forming permanent magnet ring 1910). In particular, exemplary block 2005 illustrated in FIG. 20A (e.g., a block in the inner sub-ring of outer permanent magnet ring 1910 or 310 a) may be manufactured to have dimensions as labeled in FIGS. 20B and 20C, where the dimensions are given in millimeters. Exemplary block 2105 illustrated in FIG. 21A (e.g., a block in the middle sub-ring of outer permanent magnet ring 1910 or 310 a) may be manufactured to have dimensions as labeled in FIGS. 21B and 21C, where the dimensions are also given in millimeters. Similarly, exemplary block 2205 illustrated in FIG. 22A (e.g., a block in the outer sub-ring of outer permanent magnet ring 1910 or 310 a) may be manufactured to have dimensions as labeled in FIGS. 22B and 22C, where the dimensions are also given in millimeters. It should be appreciated that the dimensions of exemplary blocks 2005, 2105 and 2205 are merely illustrative and the dimensions may be selected as desired and are not limited in this respect. Additionally, blocks may be formed using any one or combination of ferromagnetic material, as the aspects are not limited for use with any particular type of magnetic material.

FIGS. 23A and 23B illustrate different views of a portion 2315 of outer ring 2310 in quadrant(s) in the middle with respect to the yoke frame (e.g., the portion of ring 310 a in quadrant 315 b and/or 315 d illustrated in FIG. 3B), in accordance with some embodiments. That is, portion 2315 may be used for both middle quadrants, for example, in embodiments where the middle quadrants are equidistant from the yoke frame. Exemplary portion 2315 of ring 2310 includes a plurality (nine in FIGS. 23A and 23B) of stacks of three blocks each, thus forming three sub-rings of ferromagnetic blocks (e.g., blocks formed of NdFeB, SmCo, etc.). The inner sub-ring is formed of blocks (e.g., exemplary block 2305 c) having a length x, a width y and a height (or depth) z. The middle sub-ring is formed of blocks (e.g., exemplary block 2305 b) having a length X, a width y and a height z. The outer sub-ring is formed of blocks (e.g., exemplary block 2305 a) having a length X′, a width y and a height z. As shown, the blocks in the outer sub-ring have a greater length than blocks in the middle sub-ring which, in turn, have a length greater than the blocks in the inner sub-ring (i.e., x<X<X′), reducing the amount of empty space between adjacent blocks than if the blocks in all sub-rings were formed with a length x. As such more of the space in which exemplary ring 2310 is contained is occupied by field producing magnetic material, increasing the field strength in the same amount of space. It should be appreciated that the arrangement in exemplary ring 2310 is merely illustrative and other arrangements of blocks (e.g., number of stacks and number of blocks within each stack) may be used, as the aspects are not limited in this respect.

FIGS. 24A-C, 25A-C and 26A-C illustrate exemplary dimensions for blocks comprising the inner, middle and outer sub-rings of an outer permanent magnet ring (e.g., exemplary dimensions for blocks 2305 a-c forming permanent magnet ring 2310). In particular, exemplary block 2405 illustrated in FIG. 24A (e.g., a block in the inner sub-ring of outer permanent magnet ring 2310 or 310 a) may be manufactured to have dimensions as labeled in FIGS. 24B and 24C, where the dimensions are given in millimeters. Exemplary block 2505 illustrated in FIG. 25A (e.g., a block in the middle sub-ring of outer permanent magnet ring 2310 or 310 a) may be manufactured to have dimensions as labeled in FIGS. 25B and 25C, where the dimensions are also given in millimeters. Similarly, exemplary block 2605 illustrated in FIG. 26A (e.g., a block in the outer sub-ring of outer permanent magnet ring 2310 or 310 a) may be manufactured to have dimensions as labeled in FIGS. 26B and 26C, where the dimensions are also given in millimeters. It should be appreciated that the dimensions of exemplary blocks 2405, 2505 and 2605 are merely illustrative and the dimensions may be selected as desired and are not limited in this respect. Additionally, blocks may be formed using any one or combination of ferromagnetic material, as the aspects are not limited for use with any particular type of magnetic material.

FIGS. 27A and 27B illustrate different views of a portion 2715 of outer ring 2710 in a quadrant nearest the yoke frame (e.g., the portion of ring 310 a in quadrant 315 a illustrated in FIG. 3B), in accordance with some embodiments. Exemplary portion 2715 of ring 2710 includes a plurality (nine in FIGS. 27A and 27B) of stacks of three blocks each, thus forming three sub-rings of ferromagnetic blocks (e.g., blocks formed of NdFeB, SmCo, etc.). The inner sub-ring is formed of blocks (e.g., exemplary block 2705 c) having a length x, a width y and a height (or depth) z. The middle sub-ring is formed of blocks (e.g., exemplary block 2705 b) having a length X, a width y and a height z. The outer sub-ring is formed of blocks (e.g., exemplary block 2705 a) having a length X′, a width y and a height z. As shown, the blocks in the outer sub-ring have a greater length than blocks in the middle sub-ring which, in turn, have a length greater than the blocks in the inner sub-ring (i.e., x<X<X′), reducing the amount of empty space between adjacent block than if the blocks in all sub-rings were formed with a length x. As such more of the space in which exemplary ring 2710 is contained is occupied by field producing magnetic material, increasing the field strength in the same amount of space. It should be appreciated that the arrangement in exemplary ring 2710 is merely illustrative and other arrangements of blocks (e.g., number of stacks and number of blocks within each stack) may be used, as the aspects are not limited in this respect.

FIGS. 28A-C, 29A-C and 30A-C illustrate exemplary dimensions for blocks comprising the inner, middle and outer sub-rings of an outer permanent magnet ring (e.g., exemplary dimensions for blocks 2705 a-c forming permanent magnet ring 2710). In particular, exemplary block 2805 illustrated in FIG. 28A (e.g., a block in the inner sub-ring of outer permanent magnet ring 2710 or 310 a) may be manufactured to have dimensions as labeled in FIGS. 28B and 28C, where the dimensions are given in millimeters. Exemplary block 2905 illustrated in FIG. 29A (e.g., a block in the middle sub-ring of outer permanent magnet ring 2710 or 310 a) may be manufactured to have dimensions as labeled in FIGS. 29B and 29C, where the dimensions are also given in millimeters. Similarly, exemplary block 3005 illustrated in FIG. 30A (e.g., a block in the outer sub-ring of outer permanent magnet ring 2710 or 310 a) may be manufactured to have dimensions as labeled in FIGS. 30B and 30C, where the dimensions are also given in millimeters. It should be appreciated that the dimensions of exemplary blocks 2805, 2905 and 3005 are merely illustrative and the dimensions may be selected as desired and are not limited in this respect. Additionally, blocks may be formed using any one or combination of ferromagnetic material, as the aspects are not limited for use with any particular type of magnetic material.

It should be appreciated that the permanent magnet illustrated in FIG. 3A can be manufactured using any number and arrangement of permanent magnet blocks and are limited to the number, arrangement, dimensions or materials illustrated herein. The configuration of the permanent magnets will depend, at least in part, on the design characteristics of the B₀ magnet, including, but not limited to, the field strength, field of view, portability and/or cost desired for the MRI system in which the B₀ magnet is intended to operate.

As discussed above, the height or depth of the blocks used in the different quadrants may be varied to compensate for effects on the B₀ magnetic field resulting from an asymmetric yoke. For example, compare the height/depth of the exemplary blocks in portions 1915, 2315, 2715 corresponding to the far, middle and near quadrants, respectively, which are varied to compensate for effects of the yoke. In the configuration illustrated in FIG. 3A, for example, the position of frame 322 (in particular, legs 325 a and 325 b) to the permanent magnets 310 a and 310 b results in magnetic flux being drawn away from regions proximate the frame (e.g., quadrant 315 a), reducing the flux density in these regions. To address the resulting non-uniformity in the magnetic field, the height or depth of the blocks in affected regions may be varied (e.g., increased) to generate additional magnetic flux to compensate for the reduction in magnetic flux density caused by the yoke, thereby improving the homogeneity of the B₀ magnetic field within the field of view of the B₀ magnet.

The inventors have appreciated that the arrangement, dimensions and materials used for the permanent magnet blocks may be chosen to minimize the Lorentz forces produced by the B₀ coil during operation of the gradient coils. This technique can be used to reduce vibration and acoustic noise during the operation of the MRI system. According to some embodiments, the design of the permanent magnet blocks are chosen to reduce magnetic field components perpendicular to the B₀ field, i.e., parallel to the plane of the gradient coils. According to some embodiments, the outer ring of permanent magnet blocks are designed to reduce the magnetic field components responsible for vibration of the gradient coils during operation in areas outside the field of view of the MRI system, thereby reducing vibration and acoustic noise generated during operation of the MRI system.

FIGS. 31A-31C illustrate an exemplary permanent magnet disk (e.g., permanent magnet disk 310 d illustrated in FIG. 3B), in accordance with some embodiment. Permanent magnet disk 3110 is configured to be placed at center of the permanent magnet (e.g., permanent magnet 310 a and/or 310 b illustrated in FIG. 3A) to contribute to the B₀ field produced by the permanent magnet. Permanent magnet disk 3110 may be formed from any suitable ferromagnetic material (e.g., NdFeB, SmCo, etc.) and have suitable dimensions so that, when magnetized, permanent magnet disk produces a desired magnetic field. Exemplary permanent magnet disk 3110 has a diameter of 34.04 mm and a thickness of 20.08 mm, though any dimensions may be used to satisfy the design requirements of a particular B₀ magnet.

FIG. 32 illustrates a B₀ magnet 3200, in accordance with some embodiments. B₀ magnet 3200 may share design components with B₀ magnet 300 illustrated in FIG. 3A. In particular, B₀ magnet 3200 is formed by permanent magnets 3210 a and 3210 b arranged in a bi-planar geometry with a yoke 3220 coupled thereto to capture electromagnetic flux produced by the permanent magnets and transfer the flux to the opposing permanent magnet to increase the flux density between permanent magnets 3210 a and 3210 b. Each of permanent magnets 3210 a and 3210 b are formed from a plurality of concentric permanent magnets, as shown by permanent magnet 3210 b comprising an outer ring of permanent magnets 3214 a, a middle ring of permanent magnets 3214 b, an inner ring of permanent magnets 3214 c, and a permanent magnet disk 3214 d at the center. Permanent magnet 3210 a may comprise the same set of permanent magnet elements as permanent magnet 3210 b. The permanent magnet material used may be selected depending on the design requirements of the system (e.g., NdFeB, SmCo, etc. depending on the properties desired).

The permanent magnet rings are sized and arranged to produce a homogenous field of a desired strength in the central region (field of view) between permanent magnets 3210 a and 3210 b. In particular, in the exemplary embodiment illustrated in FIG. 32, each permanent magnet ring comprises a plurality of circular arc segments sized and positioned to produce a desired B₀ magnetic field, as discussed in further detail below. In a similar manner to yoke 320 illustrated in FIG. 3A, yoke 3220 is configured and arranged to capture magnetic flux generated by permanent magnets 3210 a and 3210 b and direct it to the opposing side of the B₀ magnet to increase the flux density in between permanent magnets 3210 a and 3210 b. Yoke 3220 thereby increases the field strength within the field of view of the B₀ magnet with less permanent magnet material, reducing the size, weight and cost of the B₀ magnet. Yoke 3220 also comprises a frame 3222 and plates 3224 a and 3224 b that, in a manner similar to that described above in connection with yoke 3220, captures and circulates magnetic flux generated by the permanent magnets 3210 a and via the magnetic return path of the yoke to increase the flux density in the field of view of the B₀ magnet. The structure of yoke 3220 may be similar to that described above to provide sufficient material to accommodate the magnetic flux generated by the permanent magnets and providing sufficient stability, while minimizing the amount of material used to, for example, reduce the cost and weight of the B₀ magnet.

FIG. 33 illustrates a top-down view of a permanent magnet 3310, which may, for example, be used as the design for permanent magnets 3310 a and 3310 b of B₀ magnet 3200 illustrated in FIG. 32. Permanent magnet 3310 comprises concentric rings 3310 a, 3310 b, and 3310 c, each constructed of a plurality of circular arc segments of ferromagnetic material, and a ferromagnetic disk 3310 d at the center. The direction of the frame of the yoke to which permanent magnet is attached is indicated by arrow 22. In embodiments in which the yoke is not symmetric (e.g., yoke 3220), the yoke will cause the magnetic field produced by the permanent magnets for which it captures and focuses magnetic flux to be asymmetric as well, negatively impacting the uniformity of the B₀ magnetic field. According to some embodiments, one or more dimensions of the circular arc segments are varied to compensate for the effects of the yoke on the magnetic field produced by the permanent magnet. For example, one or more dimensions of circular arc segments in the four quadrants 3315 a, 3315 b, 3315 c and 3315 d labeled in FIG. 33 may be varied to compensate for the effects of the yoke on the B₀ magnetic field, as discussed in further detail below.

FIGS. 34A and 34B illustrate different views of an inner ring 3410 (e.g., ring 3310 c illustrated in FIG. 33), in accordance with some embodiments. Exemplary ring 3410 includes a plurality (eight in exemplary ring 3410 illustrated in FIGS. 34A and 34B) of ferromagnetic circular arc segments (e.g., segments formed of NdFeB, SmCo, etc.), each spanning 45° of the ring. In exemplary ring 3410, the circular arc segments (e.g., exemplary circular arc segment 3405) are dimensioned so as to provide a ring with inner radius R1 and outer radius R2 and a height or depth z. According to some embodiments, inner ring 3410 has dimensions of R1 between 45-47 mm (e.g., 46.08 mm), R2 between 62-64 mm (e.g., 62.91 mm) and z between 23 and 24 mm (e.g., 23.46 mm). It should be appreciated that the number of circular arc segments and the dimensions thereof may be chosen as desired to produce a desired B₀ magnetic field, as the aspects are not limited in this respect.

FIGS. 35A and 35B illustrate different views of a segment 3515 that can be used to form middle ring 3510 illustrated in FIG. 35C (e.g., ring 3310 b illustrated in FIG. 33). For example, segment 3515 can be used to provide the segments in quadrants Q1-Q4 as illustrated in FIG. 35C (also, e.g., segments in quadrants 3315 a-d of ring 3310 b illustrated in FIG. 33). Exemplary portion 3515 includes a plurality of ferromagnetic circular arc segments (e.g., segments formed of NdFeB, SmCo, etc.), In FIGS. 35A-35C, two circular arc segments (e.g., exemplary circular arc segment 3505), each spanning 45°, form a quadrant of ring 3510. In exemplary portion 3515 of ring 3510, the circular arc segments are dimensioned so as to provide a ring with inner radius R1 and outer radius R2 and a height or depth z, which dimensions can be chosen for each quadrant to achieve a desired magnetic field, non-limiting examples of which are provided below.

FIGS. 36A and 36B illustrate different views of a segment 3615 that can be used to form outer ring 3610 illustrated in FIG. 36C (e.g., ring 3310 a illustrated in FIG. 33). For example, segment 3615 can be used to provide the segments in quadrants Q1-Q4 as illustrated in FIG. 36C (also, e.g., segments in quadrants 3315 a-d of ring 3310 a illustrated in FIG. 33). Exemplary portion 3615 includes a plurality of ferromagnetic circular arc segments (e.g., segments formed of NdFeB, SmCo, etc.), In FIGS. 36A-36C, five circular arc segments (e.g., exemplary circular arc segment 3605), each spanning 18° of ring 3610, form a quadrant of ring 3610. In exemplary segment 3615 of ring 3610, the circular arc segments are dimensioned so as to provide a ring with inner radius R1 and outer radius R2 and a height or depth z, which dimensions can be chosen for each quadrant to achieve a desired magnetic field.

Because a permanent B₀ magnet, once magnetized, will produce its own persistent magnetic field, power is not required to operate the permanent B₀ magnet to generate its magnetic field (e.g., exemplary B₀ magnets 300 and 3200 illustrated in FIGS. 3 and 32, respectively). As a result, a significant (often dominant) contributor to the overall power consumption of an MRI system can be eliminated, facilitating the development of an MRI system that can be powered using mains electricity (e.g., via a standard wall outlet or common large household appliance outlets), as discussed in further detail below in connection with exemplary low power MRI systems.

As discussed above, conventional power components adapted to operate a gradient coil system are generally unsuitable for use in low-field MRI due, at least in part to, expense, noise levels and/or power consumption. For example, while the cost of conventional power components used to operate gradient coils in currently available MRI systems may be acceptable given the relative insignificance compared to the total cost of a high-field MRI installation, this cost may be unacceptably high in the context of a low-field MRI system that is designed as a lower cost alternative. Thus, the cost of a power component conventionally used for high-field MRI may be disproportionately large and therefore not satisfactory for some lower cost low-field MRI systems.

Additionally, the relatively low SNR in the low-field (and particularly in the very-low and ultra-low-field regimes) renders conventional gradient coil power components unsuitable. In particular, conventional power components for driving gradient coils are generally unsuitable for low-field MRI systems because they are not designed to drive the coils with sufficiently low noise. Although the noise injected by such power components may be acceptable in the high SNR regime of high-field MRI systems, such components generally do not provide a sufficiently low level of noise to provide acceptable image quality in a low-field MRI system. For example, conventional power components may exhibit unsatisfactory variation in the output (e.g., ripple) for use in the low-field context, injecting relatively significant noise into the gradient coil system of a low-field MRI system. Additionally, conventional power components configured to drive the gradient coil system of currently available MRI systems are not designed to be power efficient, consuming large amounts of power.

The inventors have developed low-noise power component(s) suitable for driving the gradient coil system of a low-field MRI system and, according to some embodiments, have developed relatively low-noise power component(s) implemented using a relatively high efficiency linear amplifier design facilitating a low power solution, some examples of which are described in Appendix A attached hereto. Accordingly, the techniques developed by the inventors provide for a low cost, low power, low noise gradient coil system suitable for a low-field, very-low field or ultra-low field MRI system, and more particularly, for a portable MRI system that can operate using standard and/or commonly available power connections. That is, in addition to facilitating a low power MRI system, the gradient coils and gradient coil power components facilitate MRI at lower field strengths not attainable using conventional gradient coil systems due, at least in part, to the low noise operation of the gradient power components. According to some embodiments, the power electronics for powering the gradient coils of an MRI system consume less than 50 W when the system is idle and between 100-300 W when the MRI system is operating (i.e., during image acquisition), allowing for operation from standard wall power.

As discussed above, the inventors have developed electronics, including power components, that are capable of fitting within the footprint of the MRI system, thus replacing large expensive power components and other electronics shelved in separate rooms in conventional MRI systems. FIGS. 37A-37E illustrate an exemplary portable MRI system in accordance with some embodiments, illustrating the power components, power supplies, console and power distribution unit fitting within the footprint of the portable MRI system. FIG. 37A illustrates the portable low-field MRI system 3700, which is discussed in further detail below. Low field MRI system 3700 is supported by base 3750 that houses the components needed to operate the MRI system. In particular, base 3750 may be separated into compartments 3750 a, 3750 b, 3750 c housing various components of the system, as schematically illustrated in FIG. 37B and as shown in FIGS. 37C, 37D and 37E, respectfully.

Compartment 3750 a houses the gradient power amplifiers of the system and, more specifically, the three gradient power amplifiers 3765 a, 3765 b and 3765 c providing power to the X-gradient coil, the Y-gradient coil and the Z-gradient coil (generally shown as 3815), respectively. The gradient coils may be located in the magnet housing 3712 that also houses the B₀ magnet. Gradient coils 3715 may be positioned in the upper or lower portion of housing 3712, or may be positioned in both. Gradient coils 3715 may be of any suitable design, for example, any of the gradient coils described in Appendix G attached hereto.

Compartment 3750 b houses the on-system computer, console and the power distribution unit (PDU), including the backplane. Compartment 3750 c houses the power supplies, power amplifier and the radio frequency coils. Thus, the electronics used to operate the system are housed within the base on which the B₀ magnet is supported to allow the portable MRI system to be moved to desired locations. Compartment 3750 b and/or 3750 c may also house any power components needed to provide power to one or more shim or correction coils for the B₀ magnet, for example, any of the shim or correction coils described in Appendix G.

As discussed above, a factor in developing a portable MRI system is the ability to operate the MRI system in generally unshielded, partially shielded environments (e.g., outside of specially shielded rooms or encompassing cages or tents). To facilitate portable MRI that can be flexibly and widely deployed and that can be operated in different environments (e.g., an emergency room, operating room, office, clinic, etc.), the inventors have developed noise reduction systems comprising noise suppression and/or avoidance techniques for use with MRI systems in order to eliminate or mitigate unwanted electromagnetic noise, reduce its impact on the operation of the MRI systems and/or to avoid bands in the electromagnetic spectra where significant noise is exhibited.

Performance of a flexible low-field MRI systems (e.g., a generally mobile, transportable or cartable system and/or a system that can be installed in a variety of settings such as in an emergency room, office or clinic) may be particularly vulnerable to noise, such as RF interference, to which many conventional high field MRI systems are largely immune due to being installed in specialized rooms with extensive shielding. To facilitate low field MRI systems that can be flexibly and widely deployed, the inventors have developed noise reduction systems that employ one or more noise suppression techniques for use with low-field MRI systems in order to eliminate or mitigate unwanted noise or to reduce its impact on the operation of the low-field systems.

According to some embodiments, noise suppression and/or avoidance techniques are based on noise measurements obtained from the environment. The noise measurements are subsequently used to reduce the noise present in MR signals detected by the low-field MRI system (e.g., a system having a B₀ field of approximately 0.2 T or less, approximately 0.1 T or less, approximately 50 mT or less, approximately 20 mT or less, approximately 10 mT or less, etc.) during operation, either by suppressing the environmental noise, configuring the low-field MRI system to operate in a frequency band or bin having less noise, using signals obtained from multiple receive coils, or some combination therewith. Thus, the low-field MRI system compensates for noise present in whatever environment the system is deployed and can therefore operate in unshielded or partially shielded environments so that MRI is not limited to specialized shielded rooms.

Noise suppression techniques developed by the inventors are described in more detail below and it should be appreciated that the noise suppression techniques described herein may be used with any suitable low-field or high-field MRI systems deployed in virtually any facility, including portable MRI systems configured to operate outside shielded rooms and/or that have varying levels of device level shielding of the imaging region of the system. While aspects of noise suppression described herein may be particularly beneficial in the low-field context where extensive shielding may be unavailable or otherwise not provided, it should be appreciated that the techniques described herein are also suitable in the high-field context and are not limited for use with any particular type of MRI system.

Another technique for addressing the relatively low SNR characteristic of the low-field regime is to improve the homogeneity of the B₀ field by the B₀ magnet. In general, a B₀ magnet requires some level of shimming to produce a B₀ magnetic field with a profile (e.g., a B₀ magnetic field at the desired field strength and/or homogeneity) satisfactory for use in MRI. In particular, production factors such as design, manufacturing tolerances, imprecise production processes, environment, etc., give rise to field variation that produces a B₀ field having unsatisfactory profile after assembly/manufacture. For example, after production, exemplary B₀ magnets 200, 300 and/or 3200 described above may produce a B₀ field with an unsatisfactory profile (e.g., inhomogeneity in the B₀ field unsuitable for imaging) that needs to be improved or otherwise corrected, typically by shimming, to produce clinically useful images.

Shimming refers to any of various techniques for adjusting, correcting and/or improving a magnetic field, often the B₀ magnetic field of a magnetic resonance imaging device. Similarly, a shim refers to something (e.g., an object, component, device, system or combination thereof) that performs shimming (e.g., by producing a magnetic field). Techniques for facilitating more efficient and/or cost effective shimming for a B₀ magnet for MRI are described in U.S. application Ser. No. 15/466,500, titled “Methods and Apparatus for Magnetic Field Shimming” and filed on Mar. 22, 2017, which is herein incorporated by reference in its entirety, any shimming techniques of which may be used to improve the profile of any of the B₀ magnets described herein.

Using the techniques described herein, the inventors have developed portable, low power MRI systems capable of being brought to the patient, providing affordable and widely deployable MRI where it is needed. FIGS. 38A and 38B illustrate views of a portable MRI system, in accordance with some embodiments. Portable MRI system 3800 comprises a B₀ magnet 3810 formed in part by an upper magnet 3810 a and a lower magnet 3810 b having a yoke 3820 coupled thereto to increase the flux density within the imaging region. The B₀ magnet 3810 may be housed in magnet housing 3812 along with gradient coils 3815 (e.g., any of the gradient coils described in U.S. application Ser. No. 14/845,652, titled “Low Field Magnetic Resonance Imaging Methods and Apparatus” and filed on Sep. 4, 2015, which is herein incorporated by reference in its entirety). According to some embodiments, B₀ magnet 3810 comprises an electromagnet, for example, an electromagnet similar to or the same as electromagnet 210 illustrated in FIG. 2. According to some embodiments, B₀ magnet 3810 comprises a permanent magnet, for example, a permanent magnet similar to or the same as permanent magnet 300 illustrated in FIG. 3A or permanent magnet 3200 illustrated in FIG. 32.

Portable MRI system 3800 further comprises a base 3850 housing the electronics needed to operate the MRI system. For example, base 3850 may house the electronics discussed above in connection with FIGS. 37A-37E, including power components configured to operate the MRI system using mains electricity (e.g., via a connection to a standard wall outlet and/or a large appliance outlet). For example, base 3870 may house low power components, such as those described herein, enabling at least in part the portable MRI system to be powered from readily available wall outlets. Accordingly, portable MRI system 3800 can be brought to the patient and plugged into a wall outlet in the vicinity.

Portable MRI system 3800 further comprises moveable slides 3860 that can be opened and closed and positioned in a variety of configurations. Slides 3860 include electromagnetic shielding 3865, which can be made from any suitable conductive or magnetic material, to form a moveable shield to attenuate electromagnetic noise in the operating environment of the portable MRI system to shield the imaging region from at least some electromagnetic noise. As used herein, the term electromagnetic shielding refers to conductive or magnetic material configured to attenuate the electromagnetic field in a spectrum of interest and positioned or arranged to shield a space, object and/or component of interest. In the context of an MRI system, electromagnetic shielding may be used to shield electronic components (e.g., power components, cables, etc.) of the MRI system, to shield the imaging region (e.g., the field of view) of the MRI system, or both.

The degree of attenuation achieved from electromagnetic shielding depends on a number of factors including the type material used, the material thickness, the frequency spectrum for which electromagnetic shielding is desired or required, the size and shape of apertures in the electromagnetic shielding (e.g., the size of the spaces in a conductive mesh, the size of unshielded portions or gaps in the shielding, etc.) and/or the orientation of apertures relative to an incident electromagnetic field. Thus, electromagnetic shielding refers generally to any conductive or magnetic barrier that acts to attenuate at least some electromagnetic radiation and that is positioned to at least partially shield a given space, object or component by attenuating the at least some electromagnetic radiation.

It should be appreciated that the frequency spectrum for which shielding (attenuation of an electromagnetic field) is desired may differ depending on what is being shielded. For example, electromagnetic shielding for certain electronic components may be configured to attenuate different frequencies than electromagnetic shielding for the imaging region of the MRI system. Regarding the imaging region, the spectrum of interest includes frequencies which influence, impact and/or degrade the ability of the MRI system to excite and detect an MR response. In general, the spectrum of interest for the imaging region of an MRI system correspond to the frequencies about the nominal operating frequency (i.e., the Larmor frequency) at a given B₀ magnetic field strength for which the receive system is configured to or capable of detecting. This spectrum is referred to herein as the operating spectrum for the MRI system. Thus, electromagnetic shielding that provides shielding for the operating spectrum refers to conductive or magnetic material arranged or positioned to attenuate frequencies at least within the operating spectrum for at least a portion of an imaging region of the MRI system.

In portable MRI system 3800 illustrated in the moveable shields are thus configurable to provide shielding in different arrangements, which can be adjusted as needed to accommodate a patient, provide access to a patient and/or in accordance with a given imaging protocol. For example, for the imaging procedure illustrated in FIG. 39A (e.g., a brain scan), once the patient has been positioned, slides 3960 can be closed, for example, using handle 3862 to provide electromagnetic shielding 3965 around the imaging region except for the opening that accommodates the patient's upper torso. In the imaging procedure illustrated in FIG. 39B (e.g., a scan of the knee), slides 3960 may be arranged to have openings on both sides to accommodate the patient's legs. Accordingly, moveable shields allow the shielding to be configured in arrangements suitable for the imaging procedure and to facilitate positioning the patient appropriately within the imaging region.

As discussed above, a noise reduction system comprising one or more noise reduction and/or compensation techniques may also be performed to suppress at least some of the electromagnetic noise that is not blocked or sufficiently attenuated by shielding 3865. In particular, as discussed in the foregoing, the inventors have developed noise reduction systems configured to suppress, avoid and/or reject electromagnetic noise in the operating environment in which the MRI system is located. According to some embodiments, these noise suppression techniques work in conjunction with the moveable shields to facilitate operation in the various shielding configurations in which the slides may be arranged. For example, when slides 3960 are arranged as illustrated in FIG. 39B, increased levels of electromagnetic noise will likely enter the imaging region via the openings. As a result, the noise suppression component will detect increased electromagnetic noise levels and adapt the noise suppression and/or avoidance response accordingly. Due to the dynamic nature of the noise suppression and/or avoidance techniques described herein, the noise reduction system is configured to be responsive to changing noise conditions, including those resulting from different arrangements of the moveable shields. Thus, a noise reduction system in accordance with some embodiments may be configured to operate in concert with the moveable shields to suppress electromagnetic noise in the operating environment of the MRI system in any of the shielding configurations that may be utilized, including configurations that are substantially without shielding (e.g., configurations without moveable shields), as discussed in further detail below.

To ensure that the moveable shields provide shielding regardless of the arrangements in which the slides are placed, electrical gaskets may be arranged to provide continuous shielding along the periphery of the moveable shield. For example, as shown in FIG. 38B, electrical gaskets 3867 a and 3867 b (see also FIG. 45C) may be provided at the interface between slides 3860 and magnet housing to maintain to provide continuous shielding along this interface. According to some embodiments, the electrical gaskets are beryllium fingers or beryllium-copper fingers, or the like (e.g., aluminum gaskets), that maintain electrical connection between shields 3865 and ground during and after slides 3860 are moved to desired positions about the imaging region. According to some embodiments, electrical gaskets 3867 c are provided at the interface between slides 3860, as illustrated in FIG. 39B, so that continuous shielding is provided between slides in arrangements in which the slides are brought together. Accordingly, moveable slides 3860 can provide configurable shielding for the portable MRI system.

To facilitate transportation, a motorized component 3880 is provide to allow portable MRI system to be driven from location to location, for example, using a control such as a joystick or other control mechanism provided on or remote from the MRI system. In this manner, portable MRI system 3800 can be transported to the patient and maneuvered to the bedside to perform imaging, as illustrated in FIGS. 39A and 39B. As discussed above, FIG. 39A illustrates a portable MRI system 3900 that has been transported to a patient's bedside to perform a brain scan. FIG. 39B illustrates portable MRI system 3900 that has been transported to a patient's bedside to perform a scan of the patient's knee.

The portable MRI systems described herein (e.g., MRI systems illustrated in FIGS. 37-40 and 47A and 47B) may be operated from a portable electronic device, such as a notepad, tablet, smartphone, etc. For example, tablet computer 3875 may be used to operate portable MRI system to run desired imaging protocols and to view the resulting images. Tablet computer may be connected to a secure cloud to transfer images for data sharing, telemedicine and/or deep learning on the data sets. Any of the techniques of utilizing network connectivity described in Appendix F attached hereto may be utilized in connection with the portable MRI systems described herein.

FIG. 40 illustrates another example of a portable MRI system, in accordance with some embodiments. Portable MRI system 4000 may be similar in many respects to portable MRI systems illustrated in FIGS. 37-39. However, slide 4060 are constructed differently, as is shielding 4065, resulting in electromagnetic shields that are easier and less expensive to manufacture. As discussed above, a noise reduction system may be used to allow operation of a portable MRI system in unshielded rooms and with varying degrees of shielding about the imaging region on the system itself, including no, or substantially no, device-level electromagnetic shields for the imaging region.

It should be appreciated that the electromagnetic shields illustrated in FIGS. 37-40 are exemplary and providing shielding for an MRI system is not limited to the example electromagnetic shielding described herein. Electromagnetic shielding can be implemented in any suitable way using any suitable materials. For example, electromagnetic shielding may be formed using conductive meshes, fabrics, etc. that can provide a moveable “curtain” to shield the imaging region. Electromagnetic shielding may be formed using one or more conductive straps (e.g., one or more strips of conducting material) coupled to the MRI system as either a fixed, moveable or configurable component to shield the imaging region from electromagnetic interference, some examples of which are described in further detail below. Electromagnetic shielding may be provided by embedding materials in doors, slides, or any moveable or fixed portion of the housing. Electromagnetic shields may be deployed as fixed or moveable components, as the aspects are not limited in this respect.

Accordingly, aspects of the technology described herein relate to improving the performance of low-field MRI systems in environments where the presence of noise, such as RF interference, may adversely impact the performance of such systems. In some embodiments, a low-field MRI system may be configured to detect noise (e.g., environmental electromagnetic noise, internal system noise, radio frequency interference, etc.) and, in response, adapt the low-field MRI system to reduce the impact of the noise on the operation of the system. The low-field MRI system may be configured to reduce the impact of the noise by suppressing noise in the RF signal obtained by the RF receive coil, by generating RF signals that destructively interfere with noise in the environment (e.g., RF interference), by adjusting characteristics of the magnetic fields produced (e.g., adjusting the magnetic field strength of the B0 magnet) and/or received by the low-field MRI system so that the transmit/receive coils operate in a frequency band satisfactorily free from interference, or using a combination of these techniques.

According to some embodiments, noise suppression techniques described herein allow a MRI system to be operated in unshielded or partially shielded environments and/or with or without device level shielding of the image (e.g., shielding provided on the low-field MRI device itself to shield the imaging region from electromagnetic interference), at least in part by adapting noise compensation to the particular environment in which the MRI system is deployed. As a result, deployment of an MRI system is not confined to specially shielded rooms or other customized facilities and instead can be operated in a wide variety of environments.

In some embodiments, a system may be configured to obtain information about noise in the system's environment or within the system itself (e.g., RF interference) and suppress noise in the RF signal measured by the RF receive coil based, at least in part, on the obtained information. The system may be configured to obtain information about noise in the environment by using one or more auxiliary sensors. The term “auxiliary” is used to differentiate between a sensor or detector capable of detecting noise and the primary receive channel that receives MR signals for use in MRI. It should be appreciated that, in some embodiments, an auxiliary sensor may also receive one or more MR signals. For example, the low-field MRI system may comprise one or more auxiliary RF receive coils positioned proximate to the primary transmit/receive coil(s), but outside of the field of view of the B0 field, to detect RF noise without detecting MR signals emitted by a subject being imaged. The noise detected by the auxiliary RF coil(s) may be used to suppress the noise in the MR signal obtained by the primary RF coil of the MRI system.

Such an arrangement has the ability to dynamically detect and suppress RF noise to facilitate the provision of, for example, a generally transportable and/or cartable low-field MRI system that is likely to be subjected to different and/or varying levels of RF noise depending on the environment in which the low-field MRI system is operated. That is, because noise suppression is based on the current noise environment, techniques described herein provide noise suppression capability specific to the particular environment in which the system is deployed. The simplistic approach of subtracting samples of noise obtained by one or more auxiliary sensors from the signal measured by the primary receive coil(s) generally provides unsatisfactory noise suppression, even if the gain of the noise detected by the auxiliary sensor(s) is adjusted. The primary receive coil(s) and the auxiliary sensor(s) may measure different noise signals because the primary coil(s) and the auxiliary sensor(s) may be in different locations, have different orientations, and/or may have different physical characteristics (e.g., may have a different number of coil turns, may differ in size, shape, impedance, or may be a different type of sensor altogether).

Different locations and/or orientations of the primary coil(s) and the auxiliary sensor(s) may lead to differences in the characteristics of the noise signals received by the primary coil and the auxiliary sensor. Different physical characteristics between the primary coil(s) and auxiliary sensor(s) may lead to frequency-dependent differences between noise signals received by the primary coil(s) and auxiliary sensor(s). As a result, subtracting the noise signal measured by one or more auxiliary sensors from the signal measured by the primary coil(s) may not adequately suppress noise detected by the primary coil(s). Even if the noise signal measured by the auxiliary sensor(s) were scaled by a constant in an attempt to compensate for differences in the gain of the noise signals received by the primary coil(s) and auxiliary sensor(s), such compensation would not account for frequency-dependent differences in the noise signals.

Some noise suppression techniques employ a transform to suppress noise in the RF signal received by one or more primary receive coil(s) of a low-field MRI system. According to some embodiments, the transform operates to transform a noise signal received via one or multiple auxiliary sensors (e.g., one or more auxiliary RF coils and/or other types of sensors described herein) to an estimate of the noise received by the primary receive coil (or multiple primary receive coils). In some embodiments, noise suppression may comprise: (1) obtaining samples of noise by using the one or more auxiliary sensor(s); (2) obtaining samples of the MR data using the primary RF coil; (3) determining a transform; (4) transforming the noise samples using the transform; and (5) subtracting the transformed noise samples from the obtained MR data to suppress and/or eliminate noise.

The transform may be estimated from multiple (e.g., at least ten, at least 100, at least 1000, etc.) calibration measurements obtained using the auxiliary sensor(s) and primary coil(s). Multiple calibration measurements allow for estimating the transform with high accuracy. The transform may be computed in the time domain, frequency domain or a combination of both. According to some embodiments, a transform may be estimated from the plurality of calibration measurements. Multiple calibration measurements allow for estimating the amplitude and phase of the transform for a plurality of frequency bins across the frequency spectrum for which the transform is defined. For example, when processing signals using a K-point DFT (e.g., where K is an integer equal to 128, 256, 512, 1024 etc.), multiple measurements may allow for estimating the amplitude and phase of the transform for each of the K frequency bins.

In some embodiments, multiple auxiliary receive coils may be used as auxiliary sensors to suppress noise received by the primary transmit/receive coil(s) of a low-field MRI system. For example, in some embodiments, a low-field MRI system may include multiple RF coils positioned/configured to sense the MR signal emitted by the subject being imaged (e.g., multiple “primary” coils) and/or multiple coils positioned/configured to receive noise data, but to detect little or no MR signal (e.g., multiple “auxiliary” coils). Such an arrangement facilitates detection and characterization of multiple noise sources to suppress a variety of noise that may be present in a given environment. Multiple primary receive coils may also be used that factor into the noise characterization techniques described herein, as well as being used to accelerate image acquisition via parallel MR, or in other suitable ways, as discussed in further detail below.

In some embodiments, multiple auxiliary sensors may be used to perform noise compensation when there are multiple sources of noise in the environment of the low-field MRI system. For example, one or more auxiliary RF coils and/or one or more other types of sensors may be used to obtain information about the noise environment resulting from noise produced by multiple sources, which information in turn may be used to process the RF signal received by the primary receive coil(s) in order to compensate for the noise produced by multiple sources. For example, in some embodiments, a multichannel transform may be estimated from calibration measurements obtained using multiple auxiliary sensors and the primary RF coil(s), as described in more detail below. The multichannel transform may represent the relationships among the noise signals captured by the primary RF coil(s) and each of the multiple auxiliary sensors. For example, the transform may capture correlation among the noise signals received by the multiple auxiliary sensors. The transform may also capture correlation among the noise signals receive by the multiple auxiliary sensors and the noise signal received by the primary RF coil(s).

In some embodiments, multiple auxiliary sensors may be used to perform noise suppression by: (1) obtaining samples of noise by using multiple auxiliary sensors; (2) obtaining samples of the MR data using the primary RF coil(s); (3) obtaining a multichannel transform; (4) transforming the noise samples using the multichannel transform; and (5) subtracting the transformed noise samples from the obtained MR data to suppress and/or eliminate noise.

In some embodiments, the multichannel transform may be estimated from multiple (e.g., at least ten, at least 100, at least 1000, etc.) calibration measurements. According to some embodiments, multiple calibration measurements are used to estimate the amplitude and phase of the transform for a plurality of frequency bins across which the multichannel transform is defined. For example, when processing signals using a K-point DFT (e.g., where K is an integer equal to 128, 256, 512, 1024 etc.), multiple calibration measurements may allow for estimating the amplitude and phase of the multichannel transform for each of the K frequency bins.

According to some embodiments, the MR signal detected by one or more primary receive coils may also be utilized to characterize the noise to suppress or eliminate noise from the MR data. In particular, the inventors have recognized that by repeating MR data acquisitions using the same spatial encoding (e.g., by repeating a pulse sequence with the same operating parameters for the gradient coils), the “redundant” data acquired can be used to characterize the noise. For example, if a pulse sequence is repeated with the same spatial encoding multiple times, the MR data obtained should in theory be the same. Thus, the difference in the signals acquired from multiple acquisitions using the same spatial encoding can be presumed to have resulted from noise. Accordingly, multiple signals obtained from using the same spatial encoding may be phase shifted and subtracted (or added) to obtain a measure of the noise.

According to some embodiments, noise characterized in this manner can be used to compute a transform or included as a channel in a multi-channel transform, as discussed in further detail below. Alternatively, noise characterized in this manner can be used alone or in combination with other techniques to suppress noise from acquired MR signals. For example, a noise estimate obtained based on multiple MR signals obtained using the same spatial encoding may be used to suppress noise without computing a transform, as other suitable techniques may be used.

According to some embodiments, one or more sensors (e.g., one or more RF coils or other sensors capable of detecting electromagnetic fields) may be used to assess the noise background in a spectrum of interest to assess which band within the spectrum is cleanest from a noise perspective so that the transmit/receive coil(s) may be configured to operate in the identified frequency band. Accordingly, in some embodiments, a low-field MRI system may be adapted by adjusting the transmit/receive coil(s) to operate at a frequency band having less interference relative to other frequency bands in which the transmit/receive coil(s) can be configured to operate. For example, one or more auxiliary RF coils may be configured to monitor noise across multiple frequency bands over which the primary RF coil could operate and, the primary RF coil may be configured to operate at the frequency band having the least amount of noise, as determined by the measurements obtained using the auxiliary RF coils. In particular, an auxiliary RF coil may be a wideband RF coil configured to measure the noise level (e.g., noise floor) across a wide band of frequencies. Based on the noise measured across a frequency band of interest, the primary transmit/receive coil(s) (e.g., which may be a narrowband coil) may be configured to operate in a band determined to have less noise than other frequency bands. Alternatively, multiple sensors may be provided, each measuring noise levels in a respective frequency band. The primary transmit/receive coil(s) may then be configured to operate in the frequency band determined to have the least amount of noise present.

A significant source of interference for a low-field MRI system may be one or more power lines (e.g., power cords) supplying power to the low-field MRI system. Accordingly, in some embodiments, a low-field MRI system is configured to measure directly any interference due to the power line(s) and use the measurements to suppress or cancel such interference. For example, in some embodiments, a low-field MRI system may include one or more sensors coupled to a power line of the system to measure any RF signals produced or carried by the power line, and the measurements obtained by the sensor(s) may be used as part of the noise suppression techniques described herein (e.g., to further characterize the noise environment and facilitate estimation of a comprehensive transform).

In some embodiments, a low-field MRI system may include an antenna capacitively coupled to one of the power lines of the system and may be configured to use measurements obtained by the antenna to suppress noise in the RF signal received by the primary RF coil of the low-field MRI system. Such an antenna may be of any suitable type and, for example, may comprise a thin metal sheet wrapped around the power line and/or one or more capacitors coupled to the power line. A low-field MRI system may include multiple such antenna to detect noise resulting from any desired number of power lines supplying power to the system (or that otherwise impact the system) including, for example, hot lines carrying single-phase, two-phase, or three-phase power. In some instances, a low-field MRI system may include such an antenna for a ground wire. As another example, a low-field MRI system may include a sensor inductively coupled to a power line or multiple respective power lines (e.g., by use of a toroid or any other suitable method) to measure RF signals carried by the power line such that these measurements may be used to suppress noise in the RF signal measured by the primary RF coil of the low-field MRI system.

In some embodiments, a sensor's measurements of interference due to a power line may be used to suppress noise in the RF signal measured by the primary RF receive coil by estimating a transform between the primary RF receive coil and the sensor. This may be done in any suitable way and, for example, may be done using the techniques described herein for estimating a transform between the primary RF receive coil and an auxiliary RF receive coil. For example, noise characterized in this manner may be used to estimate a transform alone or may be a channel in a multi-channel transform. Noise characterized by a sensor coupled to one or more power lines may be utilized in other manners (e.g., used directly to suppress noise), as the aspects are not limited in this respect.

According to some embodiments, noise in the environment may be detected by coupling one or more sensors to one or more electromagnetic interference (EMI) shields. For example, a sensor may be connected inductively or capacitively between one or more EMI shields and ground to detect the EMI captured by the shield. Noise characterized in this manner may be used to suppress or eliminate noise from MR signals detected by the primary receive coil(s). For example, noise characterized by coupling a sensor to one or more EMI shields may be used to estimate a transform alone, or may be used as a channel in a multi-channel transform. Noise characterized by a sensor coupled to one or more EMI shields may be utilized in other manners, as the aspects are not limited in this respect.

According to some embodiments, noise from various sources are characterized using a combination of the above described techniques to determine a multi-channel transform that can be used to suppress or eliminate noise from the various noise sources. Noise measurements may be obtained during operation of the MRI system so that a multi-channel transform may be determined dynamically, allowing for noise suppression that adapts to the changing noise environment of the MRI system. However, noise in the environment may be characterized upon system start-up, when the system is moved to a different location and/or upon the occurrence of any event, and the characterized noise used to suppress and/or eliminate noise in acquired MR signals, as the techniques described herein can be applied as desired. While the noise suppression techniques may have particular benefits for a low-field MRI system, the techniques described herein are not limited for use at low-field and may be employed to suppress noise in the high-field context, as the aspects are not limited in this respect.

Referring again to FIG. 1, MRI system 100 includes controller 106 (also referred to as a console) having control electronics to send instructions to and receive information from power management system 110. Controller 106 may be configured to implement one or more pulse sequences, which are used to determine the instructions sent to power management system 110 to operate the magnetic components 120 in a desired sequence. For example, in a low-field MRI system, controller 106 may be configured to control power management system 110 to operate the magnetic components 120 in accordance with a balance steady-state free precession (bSSFP) pulse sequence, a low-field gradient echo pulse sequence, a low-field spin echo pulse sequence, a low-field inversion recovery pulse sequence, and/or any other suitable pulse sequence. Controller 106 may be implemented as hardware, software, or any suitable combination of hardware and software, as aspects of the disclosure provided herein are not limited in this respect.

In some embodiments, controller 106 may be configured to implement a pulse sequence by obtaining information about the pulse sequence from pulse sequences repository 108, which stores information for each of one or more pulse sequences. Information stored by pulse sequences repository 108 for a particular pulse sequence may be any suitable information that allows controller 106 to implement the particular pulse sequence. For example, information stored in pulse sequences repository 108 for a pulse sequence may include one or more parameters for operating magnetics components 120 in accordance with the pulse sequence (e.g., parameters for operating the RF transmit and receive coils 126, parameters for operating gradient coils 128, etc.), one or more parameters for operating power management system 110 in accordance with the pulse sequence, one or more programs comprising instructions that, when executed by controller 106, cause controller 106 to control system 100 to operate in accordance with the pulse sequence, and/or any other suitable information. Information stored in pulse sequences repository 108 may be stored on one or more non-transitory storage media.

As illustrated in FIG. 1, controller 106 also interacts with computing device 104 programmed to process received MR data. For example, computing device 104 may process received MR data to generate one or more MR images using any suitable image reconstruction process(es). Controller 106 may provide information about one or more pulse sequences to computing device 104 for the processing of data by the computing device. For example, controller 106 may provide information about one or more pulse sequences to computing device 104 and the computing device may perform an image reconstruction process based, at least in part, on the provided information.

Computing device 104 may be any electronic device that may process acquired MR data and generate one or more images of the subject being imaged. In some embodiments, computing device 104 may be a fixed electronic device such as a desktop computer, a server, a rack-mounted computer, or any other suitable fixed electronic device that may be configured to process MR data and generate one or more images of the subject being imaged. Alternatively, according to some embodiments of a low-field MRI system, computing device 104 may be a portable device such as a smart phone, a personal digital assistant, a laptop computer, a tablet computer, or any other portable device that may be configured to process MR data and generate one or images of the subject being imaged. In some embodiments, computing device 104 may comprise multiple computing devices of any suitable type, as the aspects are not limited in this respect. A user 102 may interact with workstation 104 to control aspects of the low-field MR system 100 (e.g., program the system 100 to operate in accordance with a particular pulse sequence, adjust one or more parameters of the system 100, etc.) and/or view images obtained by the low-field MR system 100.

FIG. 41A shows illustrative components of a portion of an example a MRI system that may be used for performing noise suppression, in accordance with some embodiments of the technology described herein. For example, transmit/receive system 4100 may form at least part of the transmit/receive equipment (e.g., transmit/receive coils 126, one or more controllers, etc.) of a low-field MRI system, such as any of the exemplary systems described in the above incorporated co-filed patent applications. Transmit/receive system 4100 is configured to detect MR signals emitted from excited atoms of a subject 4104 being imaged, and to characterize noise in the environment to suppress or remove the characterized noise from the detected MR signals, as discussed in further detail below.

As shown in FIG. 41A, transmit/receive system 4100 comprises a primary RF receive coil 4102 configured to measure MR signals emitted by the subject 4104 in response to an excitation pulse sequence (e.g., a pulse sequence selected from pulse sequence repository 108 and executed by controller 102). The excitation pulse sequence may be produced by primary RF receive coil 4102 and/or by one or more other transmit RF coils arranged proximate subject 4104 and configured to produce suitable MR pulse sequences when operated. Primary receive coil 4102 may be a single coil or may be a plurality of coils, which, in the latter case, may be used to perform parallel MRI. Tuning circuitry 4108 facilitates operation of primary receive coil 4102 and signals detected by RF coil(s) 4102 are provided to acquisition system 4110, which may amplify the detected signals, digitize the detected signals, and/or perform any other suitable type of processing.

Transmit/receive system 4100 also includes auxiliary sensor(s) 4106, which may include any number or type of sensor(s) configured to detect or otherwise measure noise sources in the environment and/or environmental noise produced by the MRI system itself. The noise measured by auxiliary sensor(s) 4106 may be characterized and used to suppress noise in the MR signal detected by primary RF coil(s) 4102 using techniques described in further detail below. After acquisition system 4110 processes the signals detected by RF coil(s) 4102 and auxiliary sensor(s) 4106, acquisition system 4110 may provide the processed signals to one or more other components of the MRI system for further processing (e.g., for use in forming one or more MR images of subject 4104). Acquisition system 4110 may comprise any suitable circuitry and may comprise, for example, one or more controllers and/or processors configured to control the MRI system to perform noise suppression in accordance with embodiments described herein. It should be appreciated that components illustrated in FIG. 41A may be configured to detect MR signals generated by a MRI system and, for example, the RF coils may be similar or the same as those described in the above incorporated co-field applications, or may be any other suitable type of coil.

In some embodiments, auxiliary sensor(s) 4106 may include one or more auxiliary coils 4206 configure to measure noise from one or more noise sources in the environment in which the MRI system is operating, as shown in FIG. 41B. In some instances, the auxiliary RF coil(s) 4206 may be constructed to be substantially more sensitive to ambient noise than to any noise generated by the coil itself. For example, the auxiliary RF coil 4206 may have a sufficiently large aperture and/or a number of turns such that the auxiliary coil is more sensitive to noise from the environment than to noise generated by the auxiliary coil itself. In some embodiments, auxiliary RF coil(s) 4206 may have a larger aperture and/or a greater number of turns than primary RF coil(s) 4102. However, auxiliary RF coil(s) 4206 may be the same as primary RF coil in this respect and/or may differ from primary RF coil(s) 4102 in other respects, as the techniques described herein are not limited to any particular choice of coils. For example, in some embodiments, an auxiliary sensor of a different type is used in place of an RF coil type sensor, as discussed in further detail below.

In the illustrative embodiment of FIG. 41B, auxiliary RF coil(s) 4206 is/are located a distance 4205 apart from primary RF coil 4102. The distance 4205 may be selected such that auxiliary coil(s) 4206 is/are sufficiently far away from the sample 4104 to avoid sensing MR signals emitted by the sample during imaging, but otherwise arranged as close as possible to the primary RF coil 4102 so that auxiliary coil(s) 4206 detect noise similar to the noise detected by primary coil(s) 4102. In this manner, the noise from one or more noise sources measured by auxiliary coil(s) 4206 and characterized using techniques discussed herein (e.g., by using the detected noise to calculate, at least in part, a transform that can be used to suppress and/or eliminate noise present on detected MR signals) may be representative of the noise detected by primary coil(s) 4102. It should be appreciated that auxiliary coil(s) 4206 need not be RF coils, but may be any type of sensor capable of detecting or measuring noise in the environment that may impact the performance of the MRI system, as the techniques described herein are not limited for use with any particular type of sensor.

According to some embodiments, auxiliary sensor(s) 4106 may include one or more auxiliary sensors 4306 configure to measure noise by coupling sensor(s) to one or more components of the MRI system, as schematically shown in FIG. 41C. For example, auxiliary sensors 4306 may include one or more sensors coupled to one or more components of the MRI system or otherwise arranged to detect noise produced by the MRI system. As discussed above, power cables are frequently a source of noise that can have a negative impact on the operation of the MRI system and, in particular, may produce noise that is detected by the one or more primary coils. According to some embodiments, auxiliary sensor(s) 4306 include one or more sensors coupled (e.g., capacitively or inductively) to one or more power cables of the system to detect noise produced therefrom. The detected noise may be characterized and used to suppress noise from detected MR signals, for example, by using the detected noise to produce, at least in part, a transform that characterizes noise detected by the auxiliary sensor(s) 4306, or by being directly applied to detected MR signals.

As discussed above, the low-field regime may facilitate systems that can be utilized in a wide variety of circumstances and/or that can be generally transported from one location to another. As a result, low-field MRI systems will frequently operate outside of specially shielded rooms. Thus, some low-field MRI systems may utilize partial shielding of one or more components of the system to prevent at least some EMI from reaching the shielded components. The inventors have appreciated that by coupling one or more sensors to one or more EMI shields (e.g., a Faraday cage of one or more components or the like) of the system, the noise absorbed by the one or more EMI shields can be measured, characterized and used to suppress and/or eliminate noise from detected MR signals. According to some embodiments, auxiliary sensor(s) 4306 include one or more sensors coupled between one or more EMI shields and ground to measure noise absorbed by the EMI shield that can be used to facilitate noise suppression. For example, the noise detected from the EMI shield may be used to compute, at least in part, a transform that can be utilized in suppressing and/or eliminating noise from detected MR signals. It should be appreciated that auxiliary sensor(s) 4306 may include any other type of sensor capable of detecting noise, as the aspects are not limited in this respect.

According to some embodiments, auxiliary sensor(s) 4106 include the primary coil(s) itself as illustrated in FIG. 41D, wherein the primary RF coil(s) are labeled both as primary receive coil 4102 and auxiliary sensor 4406 for the system, as the primary RF coil(s) may perform both roles in some circumstances. As discussed above, the inventors have recognized that certain pulse sequences facilitate using the signals acquired from the primary coil(s) to also suppress noise thereon. A pulse sequence refers generally to operating transmit coil(s) and gradient coil(s) in a prescribed sequence to induce an MR response. By repeating the same pulse sequence using the same spatial encoding, “redundant” MR signals can be obtained and used to estimate noise present in the MR signals.

To address the relatively low signal-to-noise ratio (SNR) of low-field MRI, pulse sequences have been utilized that repeat MR data acquisitions using the same spatial encoding (e.g., by repeating a pulse sequence with the same operating parameters to drive the gradient coils in the same manner). The MR signals obtained over multiple acquisitions are averaged to increase the SNR. For example, a balanced steady-state free precession (bSSFP) pulse sequence may be used to rapidly obtain MR data over multiple acquisitions, which acquisitions are then averaged together to increase the SNR. The term “average” is used herein to describe any type of scheme for combining the signals, including absolute average (e.g., mean), weighted average, or any other technique that can be used to increase the SNR by combining MR data from multiple acquisitions. Because the bSSFP pulse sequence does not require waiting for the net magnetization to realign with the B₀ field between successive MR data acquisitions (e.g., successive acquisitions may be obtained without needing to wait for the transverse magnetization vector to decrease to 0), multiple acquisitions may be rapidly obtained. However, any pulse sequence can be used to perform multiple acquisitions at the same location, as the aspects are not limited in this respect.

The inventors have appreciated that the MR data obtained during multiple acquisitions performed using the same spatial encoding may be used to suppress and/or eliminate noise from the detected MR signal. As discussed above, when multiple acquisitions are performed by repeating the pulse sequence with the same spatial encoding, the MR signals obtained should be the same or nearly the same and the differences can be attributed to noise. As such, phase shifting the MR signal obtained over multiple acquisitions and computing the difference between the signals provides a means for evaluating the noise corrupting the MR data. The difference may be obtained by phase shifting and either adding or subtracting the phase shifted MR signals depending on the type of pulse sequence utilized. For example, the bSSFP pulse sequence flips the polarity of the pulse sequence on subsequent acquisitions so that the difference may be computed by adding MR signals that have been appropriately shifted in phase. However, MR signals obtained using other pulse sequences that do not flip the polarity may be subtracted after being appropriately phase shifted to obtain the difference between multiple MR acquisitions. Because multiple acquisitions (e.g., 10, 20, 50, 100, 150 or more) obtained using the same spatial encoding may already be performed (and averaged) in the low-field context to achieve sufficiently large SNR, using one or more of the acquisitions to compute a noise estimate will not substantially increase acquisition times, if at all.

The computed noise (e.g., the difference between MR signals obtained over multiple acquisitions with the same spatial encoding can be used to suppress and/or eliminate the noise in the detected MR signal. According to some embodiments, the noise computed according to the above described technique may be used to, at least in part, determine a transform that can be used to suppress and/or eliminate noise in the manner discussed in further detail below. However, noise computed by determining the difference between multiple MR acquisitions can be utilized in other ways to suppress and/or eliminate noise, as the aspects are not limited in this respect. For example, noise computed based on determining the difference between multiple MR acquisitions obtained from the same location may be directly applied to detected MR signals or applied after further processing. It should be appreciated that the noise computed by comparing multiple acquisitions obtained using the same spatial encoding can be used to dynamically suppress and/or eliminate noise from the detected MR signals. In this way, noise cancellation dynamically adapts to changing noise conditions in the environment.

As discussed above, noise detected by one or more auxiliary sensors, some examples of which are described in the foregoing, may be used to characterize the noise from one or more noise sources and suppress and/or eliminate noise from detected MR signals. According to some embodiments, the noise detected by one or more auxiliary sensors is used to determine a transform that can be used to transform detected noise to an approximation of the noise detected by the one or more primary receive coils. According to some embodiments, noise detected by one or more auxiliary sensors is applied to detected MR signals to suppress noise without using a transform.

As a non-limiting example, a noise suppression component (e.g., acquisition system 4110 illustrated in FIGS. 41A-D) may suppress noise in a signal s_(pri)(t), detected by primary RF coil 4102, by using the signal s_(aux)(t), detected by auxiliary sensor 4106, and a primary-to-auxiliary sensor (PA) transform H_(PA)(ω) via the following expression:

s _(comp)(t)=s _(pri)(t)−

⁻¹ {H _(PA)(ω)S _(aux)(ω)},   (1)

where S_(aux)(ω) is the Fourier transform of s_(aux)(t),

⁻¹{ } is the inverse Fourier transform operator, and s_(comp)(t) is the noise-suppressed signal. It should be appreciated that the noise compensation calculation of Equation (1) may be implemented in any of numerous ways and, for example, may be implemented in the frequency domain or in the time domain, as the noise suppression techniques described herein are not limited in this respect. Exemplary techniques for estimating a PA transform are described in more detail below.

FIG. 42 is a flowchart of an illustrative process 4501 for performing noise suppression, in accordance with some embodiments of the technology described herein, including a detailed description of a technique for determining an exemplary transform, first with respect to a transform between an auxiliary sensor and a primary receive coil, followed by a description of a transform between multiple auxiliary sensors and a primary receive coil (multi-channel transform). It should be appreciated that a single or multi-channel transform may be computed for any number of receive coils so that noise cancellation in this respect can be performed using any number and type of auxiliary sensor and any number and type of receive coil. Process 4501 may be performed by components of any suitable MRI system and, for example, may be performed by components of MRI system 100 described with reference to FIG. 1 and the associated components illustrated in FIGS. 41A-D.

Process 4501 begins at acts 4502 and 4504, where a MRI system obtains MR data by using a primary RF coil (e.g., RF coil 4102) and obtains noise data using one or more auxiliary sensors (e.g., one or more RF coils 4206 and/or one or more other sensors 4106, 4306, 4406, etc.). As discussed above, any number of auxiliary sensors of any type may be used to characterize the noise in the environment of the MRI system. To illustrate aspects of the noise suppression techniques, the case of a primary RF coil and an auxiliary sensor is first considered. The primary RF coil and auxiliary sensor may operate to obtain MR and noise data substantially simultaneously such that the noise data acquired by the auxiliary sensor may be used to suppress noise in the MR data acquired by the primary RF coil.

The signal obtained by the primary RF coil may comprise both noise and an MR signal emitted by the sample being imaged. For example, if s_(pri)(t) represents the total signal measured by the primary RF coil, then s_(pri)(t) may be expressed as:

s _(pri)(t)=m _(pri)(t)+n _(pri)(t),

where m_(pri)(t) and n_(pri)(t) represent the MR signal and noise components of the total signal measured by the primary RF coil. Assuming that the auxiliary sensor measures a negligible amount of MR signal (due to the placement of the auxiliary sensor relative to the primary RF coil and the sample being imaged), the signal measured by the auxiliary sensor contains mostly ambient RF noise. For example, if s_(aux)(t) represents the total signal measured by the auxiliary sensor, then s_(aux)(t) may be expressed according to:

s _(aux)(t)=n _(aux)(t),

where n_(aux)(t) is noise measured by the auxiliary sensor.

As discussed above, the noise components of the signals measured by the primary RF coil and auxiliary sensor may be different (e.g., n_(pri)(t) may be different from n_(aux)(t)) due to physical differences between the primary coil and auxiliary sensor as well as differences in location and orientation. However, the inventors have appreciated that a relationship between the noise signals measured by the primary coil and the auxiliary sensor may be established since both measure noise from one or more common sources. Such a relationship may be, in some embodiments, represented by a primary to auxiliary transform. For example, the relationship may be represented by a primary to auxiliary transform H_(PA)(ω) as detailed below.

For example, in some embodiments, each of the noise signals n_(pri)(t) and n_(aux)(t) may contain noise from several independent sources including, but not limited to, noise from one or more sources in the environment of the low-field MRI system, noise generated by the primary RF coil and/or the auxiliary sensor, and noise generated by one or more other components of the MRI system (e.g., noise generated by tuning circuitry, acquisition system, power cables, etc.). Thus, for example, the noise signals n_(pri)(t) and n_(aux)(t) may be expressed as:

n _(pri)(t)=c _(pri)(t)+u _(pri)(t), and

n _(aux)(t)=c _(aux)(t)+u _(aux)(t)≅c _(aux)(t),

where c_(pri)(t) and c_(aux)(t) represent correlated noise (i.e., the signals c_(pri)(t) and c_(aux)(t) are correlated) generated by one or more common noise sources detected by the primary coil and the auxiliary sensor, respectively, and where u_(pri)(t) and u_(aux)(t) represent uncorrelated noise detected by the primary coil and auxiliary sensors, respectively (e.g., noise generated by the primary coil and auxiliary sensor themselves). As described above, in some embodiments, the auxiliary sensor may be configured such that it is more sensitive to noise from the environment than noise generated by the sensor itself. For example, the auxiliary sensor may be an auxiliary RF coil having a sufficiently large aperture and/or number of turns. As such, c_(aux)(t) may be substantially larger than u_(aux)(t) so that n_(aux)(t)≅c_(aux)(t).

Each of the noise signals c_(pri)(t) and c_(aux)(t) can be expressed in relation to the common noise source(s) through a respective measurement transform. For example, in the Fourier domain, the Fourier transforms C_(pri)(ω) and C_(aux)(ω) of noise signals c_(pri)(t) and c_(aux)(t) can be expressed as:

c _(pri)(ω)=H _(pri)(ω)C _(s)(ω)

c _(aux)(ω)=H _(aux)(ω)C _(s)(ω)

where C_(s)(ω) is the Fourier transform of a common noise source and H_(pri)(ω) and H_(aux)(ω) respectively represent the channel between the common noise source and the primary receive coil and auxiliary sensor. Combining the above equations yields:

C _(pri)(ω)=H _(PA)(ω)C _(aux)(ω),

where

${{H_{PA}(\omega)} = \frac{H_{pri}(\omega)}{H_{aux}(\omega)}},$

is the primary-to-auxiliary transform.

Returning to the discussion of process 4501, after the MR and noise signals are acquired at acts 4502 and 4504, process 4501 proceeds to act 4506, where a primary-to-auxiliary (PA) transform is obtained. In some embodiments, the PA transform may have been previously estimated so that obtaining the PA transform at act 4506 comprises accessing a representation of the PA transform (e.g., a frequency-domain or a time-domain representation of the PA transform). In other embodiments, obtaining the PA transform at act 4506 may comprise estimating and/or updating the estimate of the transform. Techniques for estimating a PA transform are described in more detail below.

Next, at act 4508, the noise data obtained at act 4504 and the PA transform obtained at act 4506 may be used to suppress or cancel noise in the MR data obtained at act 4502. This may be done using Equation (1) described above, using any equivalent formulation of Equation (1) (e.g., the entire calculation may be performed in the frequency domain), or in any other suitable way.

As described above, a primary-to-auxiliary transform may be used to suppress noise in the MR data acquired by a primary RF coil in a MRI system such as a low-field MRI system. In some embodiments, the primary-to-auxiliary transform may be estimated from calibration measurements obtained by the primary RF coil and the auxiliary sensor. This may be done in any suitable way. For example, the PA transform may be estimated from calibration measurements obtained when no MR signal is present or when the strength of the MR signal is small relative to the strength of the noise detected by the primary RF coil. As another example, the PA transform may be estimated from calibration measurements obtained when an MR signal is present (e.g., during operation of the MRI system). Any suitable number of calibration measurements may be used (e.g., at least 100, 100-1000, at least 1000, etc.). When more measurements are used, the PA transform may be estimated at a higher resolution (e.g., at more frequency values) and/or with increased fidelity with respect to the actual noise environment. The PA transform may be estimated using a least-squares estimation technique or any other suitable estimation technique, as the techniques described herein are not limited to any particular computational method.

According to some embodiments, a PA transform comprises a PA transform that is estimated from the calibration measurements. As one non-limiting example, when the signal acquired by the primary coil at times {t_(k)} does not contain any MR signal or when the strength of the MR signal is small relative to the strength of the noise detected by the primary RF coil, then s_(pri)(t_(k))=n_(pri)(t_(k)), so that the discrete Fourier transform of s_(pri)(t_(k)) is given by:

S _(pri)(ω_(k))=C _(pri)(ω_(k))+U _(pri)(ω_(k)),

where C_(pri)(ω_(k)) is the discrete Fourier transform of C_(pri)(t_(k)) and U_(pri)(ω_(k)) is the discrete Fourier transform of u_(pri)(t_(k)). Since C_(pri)(ω_(k))=H_(PA)(ω_(k))S_(ref)(ω_(k)), the discrete Fourier transform of the signal received at the primary coil may be represented as a function of the discrete Fourier transform of the signal received at the auxiliary sensor according to:

S _(pri)(ω_(k))=H _(PA)(ω_(k))S _(aux)(ω_(k))+U _(pri)(ω_(k))   (2)

Equation (2) represents a set of independent equations, one for each frequency component, ω_(k). Since both U_(pri) and H_(PA) are unknown, it may not be possible to determine H_(PA) from a single calibration measurement. If M calibration measurements (e.g., at least 10, at least 100, at least 1000 calibration measurements) are made such that multiple examples of S_(pri) and S_(aux) for each frequency component are obtained, then the PA transform can be determined despite the unknown U_(pri), via any suitable estimation technique, for example, via least squares estimation. This is so because multiple measurements may be used to average out the uncorrelated noise. Given M calibration measurements, a least squares estimator for the PA transform may be obtained by considering the following matrix equation for each frequency component ω_(k),

${\begin{bmatrix} {S_{pri}\left( \omega_{k} \right)}_{1} \\ \vdots \\ {S_{pri}\left( \omega_{k} \right)}_{M} \end{bmatrix} = {{H_{PA}\left( \omega_{k} \right)}\begin{bmatrix} {S_{aux}\left( \omega_{k} \right)}_{1} \\ \vdots \\ {S_{aux}\left( \omega_{k} \right)}_{M} \end{bmatrix}}},$

which can be solved according to:

$\begin{matrix} {{H_{PA}\left( \omega_{k} \right)} = {{{\left\{ {\begin{bmatrix} {S_{aux}\left( \omega_{k} \right)}_{1} \\ \vdots \\ {S_{aux}\left( \omega_{k} \right)}_{M} \end{bmatrix}^{T}\begin{bmatrix} {S_{aux}\left( \omega_{k} \right)}_{1} \\ \vdots \\ {S_{{aux}\;}\left( \omega_{k} \right)}_{M} \end{bmatrix}} \right\}^{- 1}\begin{bmatrix} {S_{aux}\left( \omega_{k} \right)}_{1} \\ \vdots \\ {S_{aux}\left( \omega_{k} \right)}_{M} \end{bmatrix}}^{T}\begin{bmatrix} {S_{pri}\left( \omega_{k} \right)}_{1} \\ \vdots \\ {S_{pri}\left( \omega_{k} \right)}_{M} \end{bmatrix}}.}} & \; \end{matrix}$

As may be appreciated from the foregoing, the above-described estimator uses multiple measurements (i.e., M noise signals measured by each of the primary and auxiliary coils) to estimate the value of the primary-to-auxiliary transform for multiple frequency bins. This results in significantly improved estimates of the PA transform as compared to techniques which rely on a single measurement (i.e., a single signal measured by each of the primary and auxiliary coils) to estimate the transform. Such single-measurement techniques may include scaling and time-shifting the reference signal before subtraction, which would correct for a difference in phase between the noise signal as received at a primary coil and an auxiliary coil, but (unlike the multiple measurement technique described herein) would not correct for frequency-dependent phase differences.

Another single-measurement technique may include scaling and phase adjusting the auxiliary noise signal in the frequency domain before subtracting it from the signal received at the primary coil. This could be accomplished by using the discrete Fourier transform (DFT) of the signals received by a primary coil and an auxiliary coil. The optimal scaling and phase shift can be determined by a least-squares fit across multiple frequency bins. For example, if S_(pri)(ω_(k)) is the DFT of the signal measured on the primary receive coil and S_(aux)(ω_(k)) is the DFT of the signal measured on an auxiliary coil at the same time, an average scaling and phase shift SPF for a subset of frequency bins (in the range of [k1,k2]) may be computed according to:

${SPF} = {\frac{\sum\limits_{k\; 1}^{k\; 2}{{S_{aux}\left( \omega_{k} \right)}{S_{pri}\left( \omega_{k} \right)}}}{\sum\limits_{k\; 1}^{k\; 2}{{S_{aux}\left( \omega_{k} \right)}{S_{aux}\left( \omega_{k} \right)}}}.}$

Although this single-measurement technique may be used to create a frequency-dependent correction, the method requires a tradeoff between frequency resolution of the correction and accuracy of the estimation of the scaling and phase offset. In particular, this “averaging across frequency bins of a single measurement” technique results in poor (e.g., high-variance, biased) estimation of a PA transform. In contrast, the above-described multiple measurement technique provides for an unbiased and low-variance estimator.

As described above, the inventors have appreciated that the use of multiple coils may facilitate improved MRI in a number of ways, including more robust noise detection and/or cancellation, accelerated image acquisition, etc. In embodiments where multiple primary receive coils and/or multiple auxiliary sensors are used, all of the sensors may be the same type or may be of different types. For example, in circumstances where one or more RF coils are used as sensors, none, some, or all of the coils may be shielded. As another example, the coils can have different sensitivities. When other types of sensors are used, at least some of the characteristics of the sensors and the primary receive coil(s) may necessarily be different, though some may be similar or the same.

In some embodiments, multiple auxiliary RF coils and/or primary RF coils may be used to accelerate imaging. For example, multiple RF coils used to sense noise from the same or different noise sources may also be used to perform parallel MR. In this manner, multiple RF coils may provide both noise characterization functions as well as accelerated image acquisition via their use as parallel receive coils.

In some embodiments, as described above, multiple sensors may be used to perform noise compensation in the presence of multiple noise sources. In an environment having N correlated noise sources, where N is an integer greater than one, the Fourier transforms C_(pri)(ω) and C_(aux)(ω) of noise signals c_(pri)(t) and c_(aux)(t), received by a primary coil and an auxiliary sensor can be expressed as:

C _(pri)(ω)=H _(pri,1)(ω)C ₁(ω)+H _(pri,2)(ω)C ₂(ω)+ . . . +H _(pri,N)(ω)C _(N)(ω)

C _(aux)(ω)=H _(aux,1)(ω)C ₁(ω)+H _(aux,2)(ω)C ₂(ω)+ . . . +H _(aux,N)(ω)C _(N)(ω),

where C_(j)(ω); 1≤j≤N, is a Fourier transform of a noise signal from the j^(th) noise source, H_(pri,j)(ω) is a transform between the primary coil and the j^(th) noise source, and H_(aux,j)(ω) is a transform between the auxiliary sensor and the j^(th) noise source. When the ratio H_(pri,j)(ω)/H_(aux,j)(ω) is different for one or more noise sources, it may not be possible to perform high quality noise compensation by using only a single auxiliary sensor. However, multiple auxiliary sensors may be used to perform noise compensation in this circumstance as described below.

Described below is a non-limiting example of how multiple auxiliary sensors may be used to perform noise compensation for multiple different noise sources. Without loss of generality, suppose a MR system has a primary coil and P auxiliary sensors (where P is any integer greater than or equal to 1). Further, suppose that the MR system is deployed in an environment in which there are N different noise sources (where N is an integer greater than or equal to 1). Let H_(ij)(ω) denote the transform between the i^(th) auxiliary sensor (where 1≤i≤P) and the jth noise source (where 1≤j≤N). The following set of equations relate the Fourier transforms of the signals received by the auxiliary sensors to the Fourier transforms of the noise signals produced by the noise sources:

${{\begin{bmatrix} H_{11} & \ldots & H_{1N} \\ \vdots & \ddots & \vdots \\ H_{P\; 1} & \ldots & H_{PN} \end{bmatrix}\begin{bmatrix} C_{1} \\ \vdots \\ C_{N} \end{bmatrix}} = \begin{bmatrix} C_{{aux},1} \\ \vdots \\ C_{{aux},P} \end{bmatrix}},$

where C_(aux,i); 1≤i≤P, is a Fourier transform of the signal received at the ith auxiliary sensor, C_(j)(ω); 1≤j≤N is a Fourier transform of a noise signal from the j^(th) noise source, and where the dependence of all the terms on frequency is not shown explicitly (the (ω) is suppressed for brevity), though it should be appreciated that all the terms in the above matrix equation are functions of frequency.

When the number of auxiliary sensors is greater than or equal to the number of noise sources (i.e., P>=N), the above matrix equation may be solved for the noise signals according to:

$\begin{bmatrix} C_{1} \\ \vdots \\ C_{N} \end{bmatrix} = {{{\left\{ {\begin{bmatrix} H_{11} & \ldots & H_{1N} \\ \vdots & \ddots & \vdots \\ H_{P\; 1} & \ldots & H_{PN} \end{bmatrix}^{T}\begin{bmatrix} H_{11} & \ldots & H_{1N} \\ \vdots & \ddots & \vdots \\ H_{P\; 1} & \ldots & H_{PN} \end{bmatrix}} \right\}^{- 1}\begin{bmatrix} H_{11} & \ldots & H_{1N} \\ \vdots & \ddots & \vdots \\ H_{P\; 1} & \ldots & H_{PN} \end{bmatrix}}^{T}\begin{bmatrix} C_{{aux},1} \\ \vdots \\ C_{{aux},P} \end{bmatrix}}.}$

If such a solution exists, the correlated noise measured on the primary receive coil may be expressed in relation to the measurements obtained by all of the auxiliary sensors according to:

$C_{pri} = {\begin{bmatrix} H_{{pri},1} & \ldots & H_{{pri},N} \end{bmatrix}{{\left\{ {\begin{bmatrix} H_{11} & \ldots & H_{1N} \\ \vdots & \ddots & \vdots \\ H_{P\; 1} & \ldots & H_{PN} \end{bmatrix}^{T}\begin{bmatrix} H_{11} & \ldots & H_{1N} \\ \vdots & \ddots & \vdots \\ H_{P\; 1} & \ldots & H_{PN} \end{bmatrix}} \right\}^{- 1}\begin{bmatrix} H_{11} & \ldots & H_{1N} \\ \vdots & \ddots & \vdots \\ H_{P\; 1} & \ldots & H_{PN} \end{bmatrix}}^{T}\begin{bmatrix} C_{{aux},1} \\ \vdots \\ C_{{aux},P} \end{bmatrix}}}$

A multi-channel transform H_(MPA) may be defined according to:

$H_{MPA} = {\begin{bmatrix} H_{{PA},1} & \ldots & H_{{PA},P} \end{bmatrix} = {\begin{bmatrix} H_{{pri},1} & \ldots & H_{{pri},N} \end{bmatrix}{{\left\{ {\begin{bmatrix} H_{11} & \ldots & H_{1N} \\ \vdots & \ddots & \vdots \\ H_{P\; 1} & \ldots & H_{PN} \end{bmatrix}^{T}\begin{bmatrix} H_{11} & \ldots & H_{1N} \\ \vdots & \ddots & \vdots \\ H_{P\; 1} & \ldots & H_{PN} \end{bmatrix}} \right\}^{- 1}\begin{bmatrix} H_{11} & \ldots & H_{1N} \\ \vdots & \ddots & \vdots \\ H_{P\; 1} & \ldots & H_{PN} \end{bmatrix}}^{T}.}}}$

It may then be seen that the noise measured by the primary receive coil is a linear combination of the noise signals measured on all the auxiliary coils:

$\begin{matrix} {C_{pri} = {{\begin{bmatrix} H_{{PA},1} & \ldots & H_{{PA},P} \end{bmatrix}\begin{bmatrix} C_{{aux},1} \\ \vdots \\ C_{{aux},P} \end{bmatrix}}.}} & (3) \end{matrix}$

Thus, given noise signals measured by P auxiliary sensors (e.g., the Fourier transforms of which are given by C_(aux,i) for 1≤i≤P), the above equation may be used to estimate the noise signal received at the primary receive coil (e.g., the Fourier transform of which is given by C_(pri)). In turn, the estimated noise signal may be subtracted from the overall signal measured by the primary receive coil (which signal would have both an MR signal component and a noise component) to perform noise suppression.

However, to use the above equation (3), an estimate of the multichannel primary-to-auxiliary transform H_(MPA)=[H_(PARC,1) . . . H_(PARC,P)] is needed. This may be achieved in any suitable way and, in some embodiments, may be done by making multiple measurements using the primary receive coil and the auxiliary sensors (e.g., at a time when there is no MR signal present) and using these measurements to estimate the multichannel primary-to-auxiliary transform. For example, given M measurements of noise signals at each of the P auxiliary sensors and the primary receive coil, the H_(MPA) may be estimated for each frequency component ω_(k) (where k is an index over frequency bins) using least squares estimation according to:

${\begin{bmatrix} {H_{{PA},1}\left( \omega_{k} \right)} \\ \vdots \\ {H_{{PA},P}\left( \omega_{k} \right)} \end{bmatrix} = {\left\{ {\begin{bmatrix} {S_{{aux},1}\left( \omega_{k} \right)}_{1} & \ldots & {S_{{aux},P}\left( \omega_{k} \right)}_{1} \\ \vdots & \ddots & \vdots \\ {S_{{aux},1}\left( \omega_{k} \right)}_{M} & \ldots & {S_{{aux},P}\left( \omega_{k} \right)}_{M} \end{bmatrix}^{T}\begin{bmatrix} {S_{{aux},1}\left( \omega_{k} \right)}_{1} & \ldots & {S_{{aux},P}\left( \omega_{k} \right)}_{1} \\ \vdots & \ddots & \vdots \\ {S_{{aux},1}\left( \omega_{k} \right)}_{M} & \ldots & {S_{{aux},P}\left( \omega_{k} \right)}_{M} \end{bmatrix}} \right\}^{- 1} \times \begin{bmatrix} {S_{{aux},1}\left( \omega_{k} \right)}_{1} & \ldots & {S_{{aux},P}\left( \omega_{k} \right)}_{1} \\ \vdots & \ddots & \vdots \\ {S_{{aux},1}\left( \omega_{k} \right)}_{M} & \ldots & {S_{{aux},P}\left( \omega_{k} \right)}_{M} \end{bmatrix}^{T} \times \begin{bmatrix} {S_{pri}\left( \omega_{k} \right)}_{1} \\ \vdots \\ {S_{pri}\left( \omega_{k} \right)}_{M} \end{bmatrix}}},$

where S_(aux,i)(ω_(k))_(m) represents the value of the kth frequency bin of the Fourier transform of the mth measured signal obtained by the ith auxiliary sensor, and where S_(pri)(ω_(k))_(m) represents the value of the kth frequency bin of the Fourier transform of the mth measured signal obtained by the primary receive coil. This least-squares approach provides the most complete correction when the columns of the following matrix are as orthogonal as possible to one another:

$\begin{bmatrix} H_{11} & \ldots & H_{1N} \\ \vdots & \ddots & \vdots \\ H_{P\; 1} & \ldots & H_{PN} \end{bmatrix}.$

Put another way, each auxiliary sensor may detect some or all of the different noise sources in a unique way from other auxiliary sensors. In order to correct for the presence of near field sources, multiple sensors may be placed in different locations to be more or less sensitive to some of the noise sources. In some embodiments, multiple sensors may be oriented orthogonally to one another (e.g., one sensor may be oriented in an “X” direction, another sensor may be oriented in the “Y” direction, and another sensor may be oriented in a “Z” direction). In this way, each vector of the time varying interference fields may be captured. It may also be beneficial to use one or more antennas as an auxiliary sensor to provide another orthogonal measurement.

It should be appreciated that the techniques described herein facilitate detecting noise in the environment of an MRI system using any number and/or type of sensor suitable for detecting noise produced by respective noise sources. As a result, noise from a variety of sources that may impact the performance of the MRI system may be detected and used to suppress and/or eliminate noise from MR signals detected by the MRI system during operation. Because techniques described herein operate on the particular noise environment of the MRI system, a noise reduction system employing noise suppression techniques described herein facilitate deployment of an MRI system wherever the system may be needed, eliminating the requirement that the system be installed in specially shielded rooms. The ability to dynamically adapt to changing noise environments facilitates development of MRI systems that can be deployed in generally noisy environments, including environments where noise sources may change over time. Because techniques described herein can be utilized during operation of the MRI system, the noise environment can be characterized dynamically so that it reflects the same noise environment to which the system is currently being exposed. These noise suppression and/or avoidance techniques permit the MRI system to operate in almost any environment and to dynamically adapt to and compensate for electromagnetic noise present, enabling a portable MRI system that can be transported to wherever the patient is located to perform the needed diagnostic, surgical or monitoring procedure.

A noise reduction system may include additional techniques to increase the SNR of a portable MRI system by reducing system noise, for example, by reducing inductive coupling between adjacent or neighboring RF coils in a multi-coil transmit/receive system. According to some embodiments, multiple coils can be used to both improve SNR and to facilitate noise suppression. For example, a collection of RF coils, which may be either RF signal coils (e.g., primary RF coils), RF noise coils (e.g., auxiliary RF coils) or both, may be arranged at different locations and orientations to detect a comprehensive RF field that can be characterized and compensated for using any of the noise suppression techniques discussed herein. In general, RF coils are tuned to increase coil sensitivity at a frequency of interest. However, inductive coupling between adjacent or neighboring coils (e.g., RF coils sufficiently proximate one another) degrades the sensitivity of tuned coils and significantly reduces the effectiveness of the collection of RF coils. Techniques for geometrically decoupling neighboring coils exist but place strict constraints on coil orientation and position in space, reducing the ability of the collection of RF coils to accurately detect the RF field and, as a consequence, degrading the noise rejection performance.

To address the negative impact of inductive coupling between coils, the inventors have utilized coil decoupling techniques that reduce the inductive coupling between radio frequency coils in multi-coil transmit/receive systems. For example, FIGS. 43A and 43B illustrate passive decoupling circuits configured to reduce inductive coupling between radio frequency coils in a multi-coil transmit/receive system, in accordance with some embodiments. Passive decoupling circuit 4300 a may be configured to decouple RF noise coils, for example, RF noise coils positioned outside the field of view of the MRI system that are not subjected to the relatively intense transmit B₁ field produced by the RF transmit system (i.e., one or more RF transmit coils). In this context, inductor L1 represents an RF coil configured to detect electromagnetic noise in the environment that is tuned by the circuit formed by capacitors C1, C2 and C3. Capacitors and inductors are arranged to provide a balanced differential circuit to reduce common mode noise. The tank circuit formed by L2, L3, C3 and C4 is configured to have a high impedance to ensure that the current through L1 remains small. Appropriate selection of the values for the L-C network ensures that the current passing through L1, while remaining small, has sufficient SNR for measurement at the differential output (Vout−, Vout+) of the LNA to characterize electromagnetic noise in the environment with adequate sensitivity. Equivalent impedance at the LNA input is given by:

$Z_{eq} = {\left( \frac{C_{4}}{C_{3}} \right)^{2}R}$

In the above equation, R is the equivalent losses of the primary inductance L1. Capacitor and inductor values can be chosen to attain optimal noise impedance of the LNA used for detection. FIG. 43B illustrates a passive decoupling circuit 4300 b configured to decouple RF coils that may be subjected to B₁ transmit fields. In particular, L1 may represent an RF signal coil within the field of view of the MRI system. Passive decoupling network 4300 b may be similar to passive decoupling network 4300 a in some respects, but differs in that diode D1, capacitor C3 and inductors L2 and L3 operate as a transmit/receive switch that isolates the RF coil (represented as inductor L1) from the LNA when RF signals are being transmitted by one or more RF transmit coils. Specifically, the L-C network is divided into two network portions by the transmit/receive switch to protect sensitive electronics during RF transmit cycles. During a transmit pulse, diode D1 is turned on to create a short circuit, isolating the RF signal coil from the receive electronics. The resulting L-C network provides a tank circuit with a high impedance that ensures that the current in L1 remains small. During receive cycles, diode D1 is turned off and the RF coil is connected to the LNA and tuned by the resulting balanced tank circuit configured to limit the current through L1, while allowing for sufficient signal to be detected at the differential outputs of the LNA. Thus, the RF coil is connected to a first tank circuit during transmit cycles and a second tank circuit during receive cycles of a pulse sequence. Equivalent impedance at the LNA input is then:

$Z_{eq} = {\left( \frac{L_{4}}{L_{2}} \right)^{2}R}$

The passive decoupling circuits illustrated in FIGS. 43A and 43B both operate using tuned circuits, requiring careful balancing and symmetry to operate without introducing common mode noise. Thus, the exemplary passive decoupling networks include relatively numerous and expensive components, increasing the cost and complexity of the transmit/receive and noise reduction systems. As another drawback, to change the frequency at which the transmit/receive system operates (i.e., resonates), redesign of the tuned circuits is required, thereby limiting the flexibility of the MRI system. Accordingly, the passive decoupling networks illustrated in FIGS. 43A and 43B have associated drawbacks of cost, complexity and/or inflexibility.

To address at least some of these drawbacks, the inventors have developed active decoupling techniques for reducing the coupling between neighboring RF coils in a multi-coil transmit/receive and noise reduction system. FIG. 43C illustrates an active decoupling circuit configured to reduce inductive coupling between radio frequency coils in a multi-coil transmit/receive system, in accordance with some embodiments. In decoupling circuit 43C00, inductor L1 represents an RF coil configured to measure an NMR signal. The RF coil is tuned via capacitor C3 connected in parallel to L1, and the differential outputs Vout−, Vout+ of the LNA measure the NMR signal sensed by the RF coil. The differential output of the LNA are also fed back to a second inductor L2 via resistors R1 and R2. The feedback circuit causes current flowing through inductor L2 to couple negative flux into L1 in response to signals, thus reducing the current flowing through L1 and consequently mitigating the inductive coupling effect on other nearby RF coils. L2 may be provided at a desired distance from L1 and the resistor values of R1 and R2 can be chosen so that the current through L2 achieves the desired current reduction in L1. Decoupling circuit 4300 c reduces the number of circuit elements required, thereby reducing the cost and complexity of the decoupling circuit.

The use of decoupling circuits, such as the decoupling circuits illustrated in FIGS. 43A, 43B and 43C, facilitates increasing SNR and mitigates the impact of inductive coupling on the noise rejection performance of a noise reduction system in a multi-coil transmit/receive system. It should be appreciated that other decoupling circuits may be used, as the aspects are not limited in this respect.

It should be further appreciated that a noise reduction system may include any one or more noise suppression, rejection and/or avoidance techniques described herein (e.g., one or more of dynamic noise suppression, rejection and/or avoidance techniques, one or more decoupling circuits to reduce inductive coupling, etc.) to facilitate operation of the portable MRI system in virtually any room and with virtually any device-level shielding configuration. As discussed above, conventional MRI systems operate in specially shielded rooms that provide an encompassing shielded space. As a result, MRI systems operating in specially shielded rooms have shielding for substantially 100% of the imaging region. MRI systems that operate within moveable tents or cages also have comprehensive shielding of the imaging region that endeavor to provide as close to 100% shielding of the imaging region as is practicable. To achieve portability, MRI systems according to some embodiments are configured to operate outside specially shielded rooms, tents or cages with varying levels of device-level shielding (e.g., shielding some fraction of the imaging region), including no, or substantially no, shielding of the imaging region.

The amount of electromagnetic shielding for an imaging region can be viewed as a percentage of the maximum solid angle, subtending the imaging region from its center, for which shielding is provided. Specifically, providing shielding for 100% of an imaging region means that electromagnetic shielding for at least the operating spectrum is provided over the maximum solid angle 4π steradian (sr) about the imaging region. Similarly, providing shielding for less than 75% of the imaging region means that electromagnetic shielding for at least the operating spectrum provides less than 0.75(4π) sr solid angle coverage of the imaging region, and so on. Accordingly, a specially shielded room provides shielding for substantially 100% of the imaging region for an MRI system deployed within the shielded room because shielding is provided over substantially the maximum solid angle of 4π sr. Similarly, moveable tents or cages are designed to provide shielding for as close to as 100% of the imaging region as is practicable.

The percentage of electromagnetic shielding of an imaging region of an MRI system refers to the total amount of shielding that protects the imaging region, including electromagnetic shielding provided via specially-shielded rooms, tents, cages, etc., as well as device-level electromagnetic shielding (e.g., electromagnetic shields coupled to the housing of the MRI device that provide electromagnetic shielding for the imaging region). Thus, the portable MRI systems illustrated in FIGS. 38-40, when operated outside of specially-shielded rooms or cages, have less than 100% shielding of their respective imaging regions and, in some configurations, have substantially less than 100% shielding. Providing shielding for less than 100% of the imaging region is referred to herein as providing shielding for a fraction of the imaging region, which fraction may be quantified by a specific percentage or percentage range. For example, the electromagnetic shields illustrated in FIGS. 38-40 may be adjusted to provide shielding for different fractions of the imaging region (e.g., varying degrees of shielding), such as at least between approximately 85% and approximately 50% (e.g., at approximately 85% or less, approximately 75% or less, approximately 65% or less, etc.).

It should be understood that providing shielding for a fraction of the imaging region refers to instances in which providing less than 100% shielding is intentional and/or by design (e.g., to provide access to or accommodate a patient in an MRI system operated outside a specially shielded room, tent or cage). In practice, shielding techniques are often imperfect and therefore may provide less than 100% shielding even though the intent is to provide 100% shielding for the imaging region (at least for the operating spectrum). For example, doors that are left open or ajar in specially shielded rooms, gaps in tents that go unnoticed, or openings that are not fully closed during imaging, etc., may result in less than 100% shielding even though the intent is to provide full coverage. Imperfect shielding material or construction may also result in unintentionally having less than 100% shielding. Providing shielding for a fraction of the imaging region should not be interpreted to cover these situation, as it refers to circumstances where the fractional coverage is intentional and/or by design.

FIGS. 44A-C illustrate a portable MRI system having different amounts of device-level shielding about the imaging region, in accordance with some embodiments. FIG. 44A illustrates a portable MRI system having shields 5065 that partially shield the imaging region 5095. For example, shields 5065 may be incorporated into slides 5060 that can be configured and positioned as desired to provide shielding around approximately 50% of the opening to imaging region 5095. FIG. 44B illustrates another example of a portable MRI system having shield 5165 that provides a lesser degree of shielding for imaging region 5195. For example, slide 5160 may be positioned as desired to provide shielding around approximately 25% of the opening to imaging region 5195. FIG. 44C illustrates an example of a portable MRI system without shields around the imaging region 5295, providing an alternative having substantially no device-level shielding for the open imaging region.

FIG. 44D illustrates a portable MRI system 4400 utilizing a further technique for electromagnetic shielding of the imaging region of the system, in accordance with some embodiments. In particular, in the embodiment illustrated in FIG. 44D, shielding from electromagnetic interference is achieved via one or more conductive strips connecting upper and lower portions of the B₀ magnet of the portable MRI system to form a conductive loop that counteracts at least some electromagnetic radiation that would otherwise result in interference. In the embodiment illustrated in FIG. 44D, conductive strip 4465 is electrically coupled to upper portion 4400 a, lower portion 4400 b and may also be connected to ground. In the embodiment illustrated in FIG. 44D, conductive strip 4465 is formed by a conductive braid, providing a flexible strip of material that can be coupled to the B₀ magnet with relative ease and convenience. However, conductive strip 4465 may be constructed or composed of any conductive material in any suitable form, some examples of which are described in further detail below.

The exemplary portable MRI system 4400 illustrated in FIG. 44D includes a ferromagnetic yoke 4420 that provides a magnetic path between upper magnet 4410 a and lower magnet 4410 b to capture and direct the magnetic field produced by the respective magnets to increase the magnetic flux density within the imaging region. In particular, similar to the exemplary yokes described in connection with FIGS. 2A-B, 3A and 32, yoke 4420 comprises a frame and upper and lower plates formed using a suitable ferromagnetic material or combination of materials (e.g., iron, steel, etc.). The upper and lower plates are coupled to the upper and lower magnets, respectively, to form a “magnetic circuit” that captures at least some of the magnetic field produced by the magnets and directs the captured magnetic fields via the “magnetic circuit” to increase the flux density within the imaging region of the MRI device.

The inventors have recognized that coupling conductive strip 4465 to the plates of the yoke forms a conductive loop in which current is induced by electromagnetic radiation propagating in directions through the conductive loop. This induced current will in turn produce an electromagnetic field that counteracts at least some of the electromagnetic radiation that induced the current and/or electromagnetic radiation similarly propagating through the loop. In this manner, electromagnetic interference can be reduced by the counteracting electromagnetic field produced by current induced in the conductive loop formed by the conductive strip 4465 and yoke 4220. Accordingly, the suppression of electromagnetic interference may be improved by the addition of further conductive strips forming additional conductive loops to produce counteracting electromagnetic fields when ambient electromagnetic radiation induces current in the respective conductive loop. In particular, as more conductive loops are added at different orientations, the resulting conductive loops will be responsive to more of the electromagnetic radiation present in the environment.

It should be appreciated that any number of conductive strips may be attached or affixed to the B₀ magnet to provide electromagnetic shielding. According to some embodiments, one or more additional strips 4465 connecting components of the B₀ magnet to ground may be provided about the imaging region to increase the amount of shielding arranged to protect the imaging region from electromagnetic interference (e.g., to increase the percentage of electromagnetic shielding for the imaging region). For example, a conductive strip shield may be attached every 180°, every 90°, every 45°, every 30° or at any other interval, either regularly or irregularly spaced about the imaging region, to provide a desired degree of electromagnetic shielding. It should be appreciated that any number of conductive strips may be used to achieve a desired percentage of shielding and/or to deliver a desired compromise between openness of the imaging region and comprehensiveness of the shielding for the imaging region, as discussed in further detail below.

While the conductive strip 4465 illustrated in FIG. 44D is made from a flexible material, one or more conductive strips may be formed in other ways, for example, as a rigid conductive strip, bar, rod or handle (or other suitable geometry) that electrically connects the magnets forming the B₀ magnet of the MRI system to ground. In this respect, one or more conductive strips may be arranged to serve as a handle to assist in moving the portable MRI system, to facilitate rotating the device or to assist in tilting the B₀ magnet (e.g., in conjunction with a goniometric member, examples of which are described in connection with FIGS. 45-47 below) in addition to providing electromagnetic shielding. It should be appreciated that different types of conductive strips may be used in combination (e.g., one or more flexible strips and/or one or more rigid strips) to provide electromagnetic shielding for the MRI system, as the aspects are not limited in this respect.

According to some embodiments, one or more conductive strips are configured to be removable so that conductive strips can be added and removed as desired, facilitating configurable strip shielding that provides a flexible approach to accommodate different operating environments, different imaging circumstances and/or different levels of claustrophobic affliction or unease of the patient. To facilitate configurable shielding in this respect, the housing for the magnets may include a plurality of fastening mechanisms (e.g., snaps, connectors, inserts or other mechanisms) that allow for removable attachment of conductive strips to the housing and that electrically couple the magnets to the conductive strips and to ground when a conductive strip shield is connected to the housing via a respective fastening mechanism. Fastening mechanisms may be arranged at any desired location and at any number of locations to provide flexibility in where and how many conductive strips may be attached to the device. Additionally, the fastening mechanisms themselves may be made to be moveable so that one or more conductive strips coupled to the system via the fastening mechanisms may be adjusted (e.g., rotated about the imaging region). In this manner, conductive strips may be added, removed and/or their positions adjusted as needed to provide a desired shielding configuration in a desired amount (e.g., to provide shielding for a desired percentage of the imaging region).

Providing a plurality of fastening mechanisms that allow removable strips to be attached and removed at a number of locations about the imaging region allows the imaging region to remain essentially open while positioning a patient within the imaging region. After the patient has been positioned within the imaging region, a desired number of conductive strips may be attached to the B₀ magnet via the plurality of fastening mechanisms to achieve a desired degree of shielding, to address the electromagnetic environment in which the MRI system is operating, to facilitate a particular imaging protocol and/or to accommodate a patient who may be susceptible to claustrophobia (e.g., conductive strips may be added only while the patient remains comfortable with the openness of the MRI system). Accordingly, strip shielding techniques may provide a flexible, configurable approach to electromagnetic shielding, facilitating the ability to deploy portable MRI systems in a variety of environments and for a variety of applications and circumstances.

There are a number of benefits to reducing the shielding provided around the imaging region (e.g., using any of the shielding techniques described herein), including a reduction in cost and complexity of the system and improvements in accessibility to the imaging region both with respect to positioning a patient for imaging, as well as increased accessibility for medical personnel who may need to perform other tasks requiring access to the patient while the patient remains positioned within the system. In addition, reducing the shielding around the imaging region maximizes the openness of the MRI system to improve the experience of patients who are susceptible to feelings of claustrophobia. In this manner, the applicability of portable MRI may be further increased from a cost and/or flexibility perspective.

According to some embodiments, device-level shields are removable such that the amount of shielding provided may be selected in view of the particular circumstances, such as the required accessibility to the patient and/or imaging region for a given procedure, the severity of a patient's claustrophobia, the particular noise environment, etc. For example, slides carrying shields may be configured to be attached and removed from the B₀ magnet, allowing for a portable MRI device to be selectively and dynamically configured as desired (e.g., to allow a portable MRI system to be configured with the amount of shielding and accessibility illustrated in FIGS. 40 and 44A-C, which have three, two, one and zero slides/shields attached, respectively). In this manner, a portable MRI device can take advantage of the shielding and accessibility aspects of the different possible configurations, allowing the portable MRI to be optimized in this respect for given procedures and/or particular patients. According to other embodiments, the number of shields or the amount of shielding for a given portable MRI system may be fixed, which may allow for reductions in cost and complexity, but may also decrease the flexibility of the system from a shielding/accessibility perspective.

As discussed above, the inventors have developed noise reduction systems that allow a portable MRI device to operate in different noise environments (e.g., in unshielded or partially shielded rooms) and to operate with varying amounts of device-level shielding. A portable MRI system may include a noise reduction system that includes any one or combination of the noise suppression, avoidance and/or reduction techniques described herein, as the aspects are not limited in this respect. For example, a noise reduction system may employ one or more of the noise suppression and/or avoidance techniques described herein, allowing for dynamic noise suppression and/or avoidance that compensates for a given noise environment and/or that works in concert with the variable amounts of device-level shielding provided by portable MRI systems having configurable shields (e.g., the portable MRI systems illustrated in FIGS. 40 and 44A-C), including no or substantially no device-level shielding about the imaging region. A noise reduction system may also include coil decoupling networks to reduce the noise resulting from inductive coupling between radio frequency coils in multi-coil transmit/receive systems at any level of shielding provided. It should be appreciated that a noise reduction system may include any one or combination of techniques described herein, as the aspects are not limited in this respect.

According to some embodiments, a portable MRI system (e.g., any of the portable MRI systems illustrated in FIGS. 37-40 and 44A-D) is configured to operate using mains electricity (e.g., single-phase electricity provided at standard wall outlets) via a power connection 3870 (see e.g., FIG. 38B). According to some embodiments, a portable MRI system comprises a power connection configured to connect to a single-phase outlet providing approximately between 110 and 120 volts and rated at 15, 20 or 30 amperes, and wherein the power system is capable of providing the power to operate the portable MRI system from power provided by the single-phase outlet. According to some embodiments, a portable MRI system comprises a power connection configured to connect to a single-phase outlet providing approximately between 220 and 240 volts and rated at 15, 20 or 30 amperes, and wherein the power system is capable of providing the power to operate the magnetic resonance imaging system from power provided by the single-phase outlet. According to some embodiments, a portable MRI system is configured using the low power techniques described herein to use an average of less than 3 kilowatts during image acquisition. According to some embodiments, a portable MRI system is configured using the low power techniques described herein to use an average of less than 2 kilowatts during image acquisition. According to some embodiments, a portable MRI system is configured using the low power techniques described herein to use an average of less than 1 kilowatt during image acquisition. For example, a low power MRI system employing a permanent B₀ magnet and low power components described herein may operate at 1 kilowatt or less, such as at 750 watts or less.

As shown in FIGS. 38-40, the portable MRI is configured so that the B₀ magnet can be tilted at a desired incline. In many instances, a patient may not be able to lie flat, for example, due to risks associated with increased hydrostatic pressure in the brain. The inventors have developed a portable MRI device having a positioning mechanism that allows the B₀ magnet to be rotated, for example, about its center of mass. Thus, if a patient or a particular portion of a patient's anatomy needs to be supported at an incline, the positioning mechanism can be engaged to rotate or tilt the B₀ magnet to achieve the desired incline. According to some embodiments, the positioning mechanism can be manually engaged to rotate or tilt the B₀ magnet by hand, facilitating quick and easy configuring of the MRI system at the desired incline.

FIGS. 45A-45D illustrate different views of a positioning mechanism that employs a positioning goniometer or goniometric stage 4590 that allows the magnet to be rotated about a fixed axis (e.g., the axis through or near the center of mass of the B₀ magnet). As illustrated in FIG. 45A, a goniometric stage 4590 is rotatably coupled to the bottom of a lower portion of the B₀ magnet to allow the B₀ magnet to be rotated about its center of mass 4591, as shown by direction arrows 4593 in FIG. 45A. Goniometric stage 4590 includes a number of holes or bores 4595 configured to accommodate a locking member (e.g., a locking pin) that locks the mechanism in place at a desired angle, as discussed in further detail below. Rotating the B₀ magnet via goniometric stage 4590 effects a tilt that provides an inclined supporting surface for the patient anatomy being imaged, as illustrated in FIG. 39A.

FIG. 45B illustrates a side view of the B₀ magnet and the goniometric stage 4590. Goniometric stage 4590 includes a release mechanism 4594 that engages and disengages a locking pin 4596 from holes provided on a fixed or stationary member of the goniometric stage 4590 (e.g., holes 4595 illustrated in FIG. 45A). To rotate the B₀ magnet, release mechanism 4594 is pressed in an upward direction to disengage the locking pin 4596 from the hole in which it is currently positioned. For example, handle 4592 allows a user to place a hand on the handle and squeeze release mechanism 4594 towards the handle to release the locking pin 4596, as discussed in further detail below in connection with FIGS. 46A and 46B. With the locking pin 4596 disengaged, the B₀ magnet may then be rotated or tilted to the desired incline using handle 4592. Once the B₀ magnet has been rotated to the desired angle, release mechanism 4594 may be released so that the locking pin 4596 engages with a corresponding hole at the new position, locking the mechanism in place at the desired angle. FIGS. 45C and 45D illustrate the goniometric stage 4590 coupled to the bottom side of the lower magnet apparatus 4510 and base 4550.

FIGS. 46A and 46B illustrate a closer view of the exemplary goniometric stage 4590 discussed in the foregoing. As shown, release mechanism 4594 is rotatably coupled to moveable stage component 4590 a via axle 4599. When force is applied to release mechanism 4594 in the direction shown by arrow 4597 (e.g., by gripping handle 4592 and release mechanism 4594 with a hand and squeezing or lifting the release mechanism towards the handle), release mechanism 4594 rotates about axle 4599 and raises portion 4598 to lift locking pin 4596 out of hole 4595 c in which it is currently positioned. When locking pin 4596 is lifted from the hole as shown by the phantom lines in FIG. 46B, moveable stage component 4590 a is released from its locked position and allowed it to slide within stationary stage component 4590 b. When the moveable stage component 4590 a is moved to its desired location, release mechanism 4594 can be release to lock moveable stage component 4590 a into the desired position. For example, a spring mechanism may be coupled to the locking pin so that when release mechanism 4594 is released, the spring force causes the locking pin 4596 to return to its locked position. While exemplary goniometric stage 4590 includes four holes (e.g., holes 4595 a, 4595 b, 4595 c and 4595 d), any number of holes at any location may be provided to provide a desired granularity to the angles at which the B₀ magnet can be positioned, as the aspects are not limited in this respect. It should be appreciated that goniometric stage 4590 allows the B₀ magnet to be rotated without movement of the center of mass, permitting the magnet to be rotated by hand. However, other mechanisms that rotate the center of mass may also be used, as the aspects are not limited in this respect.

FIGS. 47A and 47B illustrate a low power, portable low-field MRI system, in accordance with some embodiments. Portable MRI system 4700 comprises a B₀ magnet 4705 including at least one first permanent magnet 4710 a and at least one second permanent magnet 4710 b magnetically coupled to one another by a ferromagnetic yoke 4720 configured to capture and channel magnetic flux to increase the magnetic flux density within the imaging region (field of view) of the MRI system. Permanent magnets 4710 a and 4710 b may be constructed using any suitable technique, including any of the techniques described herein (e.g., using any of the techniques, designs and/or materials described in connection with B₀ magnet 300 illustrated in FIG. 3A and/or B₀ magnet 3200 illustrated in FIG. 32 and described in the accompanying description thereof). Yoke 4720 may also be constructed using any of the techniques described herein (e.g., using any of the techniques, designs and/or materials described in connection with yokes 320 and 3220 illustrated in FIG. 3A and FIG. 32 and described in the accompanying description thereof). It should be appreciated that, in some embodiments, B₀ magnet 4705 may be formed using electromagnets using any of the electromagnet techniques described herein (e.g., using any of the techniques, designs and/or materials described in connection with B₀ magnet 200 illustrated in FIGS. 2A and 2B and described in the accompanying description thereof).

B₀ magnet 4705 may be coupled to or otherwise attached or mounted to base 4750 by a positioning mechanism 4790, such as the goniometric stage described in connection with FIGS. 45A-D and 46A-B, so that the B₀ magnet can be tilted (e.g., rotated about its center of mass) to provide an incline to accommodate a patient's anatomy as needed. In FIG. 47A, the B₀ magnet is shown level without an incline and, in FIG. 47B, the B₀ magnet is shown after undergoing a rotation to incline the surface supporting the patient's anatomy being scanned. Positioning mechanism 4790 may be fixed to one or more load bearing structure of base 4750 arranged to support the weight of B₀ magnet 4700.

In addition to providing the load bearing structures for supporting the B₀ magnet, base 4750 also includes an interior space configured to house the electronics 4770 needed to operate the portable MRI system 4700. For example, base 4750 may house the power components to operate the gradient coils (e.g., X, Y and Z) and the RF transmit/receive coils. According to some embodiments, the power components providing power to operate the gradient coils comprise three gradient power components, for example, power components using a linear amplifier as described in Appendix A. According to some embodiments, the power electronics for powering the gradient coils of an MRI system consume less than 50 W when the system is idle and between 100-300 W when the MRI system is operating (i.e., during image acquisition). Base 4750 also house the RF coil amplifiers (i.e., power amplifiers to operate the transmit/receive coils of the system), power supplies, console, power distribution unit and other electronics needed to operate the MRI system.

According to some embodiments, the electronics 4770 needed to operate portable MRI system 4700 consume less than 1 kW of power and, in some embodiments, less than 750 W of power (e.g., MRI systems utilizing a permanent B₀ magnet solution). However, systems that consume greater power may also be utilized as well, as the aspects are not limited in this respect. Exemplary portable MRI system 4700 illustrated in FIGS. 47A and 47B may be powered via a single power connection 4775 configured to connect to a source of mains electricity, such as an outlet providing single-phase power (e.g., a standard or large appliance outlet). Accordingly, the portable MRI system can be plugged into a single available power outlet and operated therefrom.

Portable MRI system 4700 illustrated in FIGS. 47A and 47B also comprises a conveyance mechanism 4780 that allows the portable MRI system to be transported to different locations. The conveyance mechanism may comprise one or more components configured to facilitate movement of the portable MRI system, for example, to a location at which MRI is needed. According to some embodiments, conveyance mechanism comprises a motor 4786 coupled to drive wheels 4784. In this manner, conveyance mechanism 4780 provides motorized assistance in transporting MRI system 4700 to desired locations. Conveyance mechanism 4780 may also include a plurality of castors 4782 to assist with support and stability as well as facilitating transport.

According to some embodiments, conveyance mechanism 4780 includes motorized assistance controlled via a joystick to guide the portable MRI system during transportation to desired locations. According to some embodiments, the conveyance mechanism comprises power assist means configured to detect when force is applied to the MRI system and to, in response, engage the conveyance mechanism to provide motorized assistance in the direction of the detected force. For example, rail 4755 of base 4750 illustrated in FIGS. 47A and 47B may be configured to detect when force is applied to the rail (e.g., by personnel pushing on the rail) and engage the conveyance mechanism to provide motorized assistance to drive the wheels in the direction of the applied force. As a result, a user can guide the portable MRI system with the assistance of the conveyance mechanism that responds to the direction of force applied by the user. The power assist mechanism may also provide a safety mechanism for collisions. In particular, the force of contact with another object (e.g., a wall, bed or other structure) will also be detected and the conveyance mechanism will react accordingly with a motorized locomotion response away from the object. According to some embodiments, motorized assistance may be eliminated and the portable MRI system may be transported by having personnel move the system to desired locations using manual force.

Portable MRI system 4700 includes slides 4760 that provide electromagnetic shielding to the imaging region of the system. Slides 4760 may be transparent or translucent (e.g., slides 3860 in FIGS. 38A and 38B) to preserve the feeling of openness of the MRI system to assist patients who may experience claustrophobia during conventional MRI performed within a closed bore. Slides 4760 may also be perforated to allow air flow to increase the sense of openness and/or to dissipate acoustic noise generated by the MRI system during operation. The slides may have shielding 4765 incorporated therein to block electromagnetic noise from reaching the imaging region. According to some embodiments, slides 4760 may also be formed by a conductive mesh (e.g., slides 4060 illustrated in FIG. 40) providing shielding 4765 to the imaging region and promoting a sense of openness for the system. As discussed above, the slides provide electromagnetic shielding that is moveable to allow a patient to be positioned within the system, permitting adjustment by personnel once a patient is positioned or during acquisition, and/or enabling a surgeon to gain access to the patient, etc. Thus, the moveable shielding facilitates flexibility that allows the portable MRI system to not only be utilized in unshielded rooms, but enables procedures to be performed that are otherwise unavailable.

To further suppress and/or avoid electromagnetic noise, portable MRI system 4700 may be equipped with a noise reduction system using one or more of the noise suppression and/or avoidance techniques described herein. As also discussed above, the noise suppression/avoidance components dynamically adapts the noise suppression/cancellation response in concert with the shielding configuration of a given arrangement of the moveable shields. Thus, portable low field MRI system 4700 can be transported to the patient and/or to a desired location and operated outside specially shielded rooms (e.g., in an emergency room, operating room, NICU, general practitioner's office, clinic and/or brought bedside directly to the patient wherever located, allowing for MRI to be performed when and where it is needed.

Having thus described several aspects and embodiments of the technology set forth in the disclosure, it is to be appreciated that various alterations, modifications, and improvements will readily occur to those skilled in the art. Such alterations, modifications, and improvements are intended to be within the spirit and scope of the technology described herein. For example, those of ordinary skill in the art will readily envision a variety of other means and/or structures for performing the function and/or obtaining the results and/or one or more of the advantages described herein, and each of such variations and/or modifications is deemed to be within the scope of the embodiments described herein. Those skilled in the art will recognize, or be able to ascertain using no more than routine experimentation, many equivalents to the specific embodiments described herein. It is, therefore, to be understood that the foregoing embodiments are presented by way of example only and that, within the scope of the appended claims and equivalents thereto, inventive embodiments may be practiced otherwise than as specifically described. In addition, any combination of two or more features, systems, articles, materials, kits, and/or methods described herein, if such features, systems, articles, materials, kits, and/or methods are not mutually inconsistent, is included within the scope of the present disclosure.

The above-described embodiments can be implemented in any of numerous ways. One or more aspects and embodiments of the present disclosure involving the performance of processes or methods may utilize program instructions executable by a device (e.g., a computer, a processor, or other device) to perform, or control performance of, the processes or methods. In this respect, various inventive concepts may be embodied as a computer readable storage medium (or multiple computer readable storage media) (e.g., a computer memory, one or more floppy discs, compact discs, optical discs, magnetic tapes, flash memories, circuit configurations in Field Programmable Gate Arrays or other semiconductor devices, or other tangible computer storage medium) encoded with one or more programs that, when executed on one or more computers or other processors, perform methods that implement one or more of the various embodiments described above. The computer readable medium or media can be transportable, such that the program or programs stored thereon can be loaded onto one or more different computers or other processors to implement various ones of the aspects described above. In some embodiments, computer readable media may be non-transitory media.

The terms “program” or “software” are used herein in a generic sense to refer to any type of computer code or set of computer-executable instructions that can be employed to program a computer or other processor to implement various aspects as described above. Additionally, it should be appreciated that according to one aspect, one or more computer programs that when executed perform methods of the present disclosure need not reside on a single computer or processor, but may be distributed in a modular fashion among a number of different computers or processors to implement various aspects of the present disclosure.

Computer-executable instructions may be in many forms, such as program modules, executed by one or more computers or other devices. Generally, program modules include routines, programs, objects, components, data structures, etc. that perform particular tasks or implement particular abstract data types. Typically the functionality of the program modules may be combined or distributed as desired in various embodiments.

Also, data structures may be stored in computer-readable media in any suitable form. For simplicity of illustration, data structures may be shown to have fields that are related through location in the data structure. Such relationships may likewise be achieved by assigning storage for the fields with locations in a computer-readable medium that convey relationship between the fields. However, any suitable mechanism may be used to establish a relationship between information in fields of a data structure, including through the use of pointers, tags or other mechanisms that establish relationship between data elements.

The above-described embodiments of the present invention can be implemented in any of numerous ways. For example, the embodiments may be implemented using hardware, software or a combination thereof. When implemented in software, the software code can be executed on any suitable processor or collection of processors, whether provided in a single computer or distributed among multiple computers. It should be appreciated that any component or collection of components that perform the functions described above can be generically considered as a controller that controls the above-discussed function. A controller can be implemented in numerous ways, such as with dedicated hardware, or with general purpose hardware (e.g., one or more processor) that is programmed using microcode or software to perform the functions recited above, and may be implemented in a combination of ways when the controller corresponds to multiple components of a system.

Further, it should be appreciated that a computer may be embodied in any of a number of forms, such as a rack-mounted computer, a desktop computer, a laptop computer, or a tablet computer, as non-limiting examples. Additionally, a computer may be embedded in a device not generally regarded as a computer but with suitable processing capabilities, including a Personal Digital Assistant (PDA), a smartphone or any other suitable portable or fixed electronic device.

Also, a computer may have one or more input and output devices. These devices can be used, among other things, to present a user interface. Examples of output devices that can be used to provide a user interface include printers or display screens for visual presentation of output and speakers or other sound generating devices for audible presentation of output. Examples of input devices that can be used for a user interface include keyboards, and pointing devices, such as mice, touch pads, and digitizing tablets. As another example, a computer may receive input information through speech recognition or in other audible formats.

Such computers may be interconnected by one or more networks in any suitable form, including a local area network or a wide area network, such as an enterprise network, and intelligent network (IN) or the Internet. Such networks may be based on any suitable technology and may operate according to any suitable protocol and may include wireless networks, wired networks or fiber optic networks.

Also, as described, some aspects may be embodied as one or more methods. The acts performed as part of the method may be ordered in any suitable way. Accordingly, embodiments may be constructed in which acts are performed in an order different than illustrated, which may include performing some acts simultaneously, even though shown as sequential acts in illustrative embodiments.

All definitions, as defined and used herein, should be understood to control over dictionary definitions, definitions in documents incorporated by reference, and/or ordinary meanings of the defined terms.

The indefinite articles “a” and “an,” as used herein in the specification and in the claims, unless clearly indicated to the contrary, should be understood to mean “at least one.”

The phrase “and/or,” as used herein in the specification and in the claims, should be understood to mean “either or both” of the elements so conjoined, i.e., elements that are conjunctively present in some cases and disjunctively present in other cases. Multiple elements listed with “and/or” should be construed in the same fashion, i.e., “one or more” of the elements so conjoined. Other elements may optionally be present other than the elements specifically identified by the “and/or” clause, whether related or unrelated to those elements specifically identified. Thus, as a non-limiting example, a reference to “A and/or B”, when used in conjunction with open-ended language such as “comprising” can refer, in one embodiment, to A only (optionally including elements other than B); in another embodiment, to B only (optionally including elements other than A); in yet another embodiment, to both A and B (optionally including other elements); etc.

As used herein in the specification and in the claims, the phrase “at least one,” in reference to a list of one or more elements, should be understood to mean at least one element selected from any one or more of the elements in the list of elements, but not necessarily including at least one of each and every element specifically listed within the list of elements and not excluding any combinations of elements in the list of elements. This definition also allows that elements may optionally be present other than the elements specifically identified within the list of elements to which the phrase “at least one” refers, whether related or unrelated to those elements specifically identified. Thus, as a non-limiting example, “at least one of A and B” (or, equivalently, “at least one of A or B,” or, equivalently “at least one of A and/or B”) can refer, in one embodiment, to at least one, optionally including more than one, A, with no B present (and optionally including elements other than B); in another embodiment, to at least one, optionally including more than one, B, with no A present (and optionally including elements other than A); in yet another embodiment, to at least one, optionally including more than one, A, and at least one, optionally including more than one, B (and optionally including other elements); etc.

Also, the phraseology and terminology used herein is for the purpose of description and should not be regarded as limiting. The use of “including,” “comprising,” or “having,” “containing,” “involving,” and variations thereof herein, is meant to encompass the items listed thereafter and equivalents thereof as well as additional items.

In the claims, as well as in the specification above, all transitional phrases such as “comprising,” “including,” “carrying,” “having,” “containing,” “involving,” “holding,” “composed of,” and the like are to be understood to be open-ended, i.e., to mean including but not limited to. Only the transitional phrases “consisting of” and “consisting essentially of” shall be closed or semi-closed transitional phrases, respectively. 

What is claimed is:
 1. A portable magnetic resonance imaging system comprising: a B₀ magnet configured to produce a B₀ magnetic field for an imaging region of the magnetic resonance imaging system; a noise reduction system configured to detect and suppress at least some electromagnetic noise in an operating environment of the portable magnetic resonance imaging system; and electromagnetic shielding provided to attenuate at least some of the electromagnetic noise in the operating environment of the portable magnetic resonance imaging system, the electromagnetic shielding arranged to shield a fraction of the imaging region of the portable magnetic resonance imaging system.
 2. The portable magnetic resonance imaging system of claim 1, wherein the electromagnetic shielding is arranged to shield less than seventy-five percent of the imaging region of the magnetic resonance imaging system.
 3. The portable magnetic resonance imaging system of claim 2, wherein the electromagnetic shielding is arranged to shield less than fifty percent of the imaging region of the magnetic resonance imaging system.
 4. The portable magnetic resonance imaging system of claim 3, wherein the electromagnetic shielding is arranged to shield less than twenty-five percent of the imaging region of the magnetic resonance imaging system.
 5. The portable magnetic resonance imaging system of claim 4, wherein the electromagnetic shielding is arranged to shield less than ten percent of the imaging region of the magnetic resonance imaging system.
 6. The portable magnetic resonance imaging system of claim 1, further comprising a housing for at least the B₀ magnet, wherein the electromagnetic shielding comprises at least one electromagnetic shield structure adjustably coupled to the housing to shield the imaging region.
 7. The portable magnetic resonance imaging system of claim 6, wherein the at least one electromagnetic shield structure is adjustable to shield in a range at least between seventy-five percent and fifty percent of the imaging region.
 8. The portable magnetic resonance imaging system of claim 6, wherein the at least one electromagnetic shield structure is adjustable to shield in a range in a range at least between seventy-five percent and twenty-five percent of the imaging region.
 9. The portable magnetic resonance imaging system of claim 6, wherein the at least one electromagnetic shield structure is adjustable to shield in a range at least between fifty percent and twenty-five percent of the imaging region.
 10. The portable magnetic resonance imaging system of claim 6, wherein the at least one electromagnetic shield structure comprises a plurality of moveable shields coupled to the housing.
 11. The portable magnetic resonance imaging system of claim 10, wherein the plurality of moveable shields are capable of being moved independently of one another to allow for the amount of shielding of the imaging region to be varied.
 12. The portable magnetic resonance imaging system of claim 6, wherein the noise reduction system is configured to dynamically suppress and/or compensate for at least some of the electromagnetic noise in response to different arrangement of the at least one electromagnetic shield structure.
 13. The portable magnetic resonance imaging system of claim 1, wherein the at least one radio frequency coil comprises a plurality of radio frequency coils, and wherein the portable magnetic resonance imaging system comprises at least one decoupling circuit coupled to at least one of the plurality of radio frequency coils to reduce an amount of current passing through the at least one of the plurality of radio frequency coils.
 14. The portable magnetic resonance imaging system of claim 13, wherein the at least one decoupling circuit comprises a tank circuit that limits current that flows through the at least one of the plurality of radio frequency coils.
 15. The portable magnetic resonance imaging system of claim 14, wherein the at least one decoupling circuit comprises: a first tank circuit configured to limit current that flows through the at least one of the plurality of radio frequency coils; a second tank circuit configured to limit current that flows through the at least one of the plurality of radio frequency coils; and a transmit/receive switch that couples the first tank circuit to the at least one of the plurality of radio frequency coils during transmit cycles and couples the second tank circuit to the at least one of the plurality of radio frequency coils during receive cycles.
 16. The portable magnetic resonance imaging system of claim 15, wherein the first tank circuit and the second tank circuit share at least one circuit component.
 17. The portable magnetic resonance imaging system of claim 13, wherein the at least one decoupling circuit comprises an inductor coupled in parallel with the at least one of the plurality of radio frequency coils via a feedback circuit to reduce an amount of current flowing through the at least one of the plurality of radio frequency coils.
 18. The portable magnetic resonance imaging system of claim 13, wherein the at least one radio frequency coil comprises at least one primary radio frequency coil and at least one auxiliary radio frequency coil, and wherein the at least one decoupling circuit comprises a first decoupling circuit coupled to the at least one primary frequency coil and a second decoupling circuit coupled to the at least one auxiliary coil.
 19. The portable magnetic resonance imaging system of claim 18, wherein the first decoupling circuit and the second decoupling circuit are of a different type.
 20. The portable magnetic resonance imaging system of claim 18, wherein the first decoupling circuit and the second decoupling circuit are of a same type.
 21. The portable magnetic resonance imaging system of claim 1, wherein the B₀ magnet is configured to produce a B₀ field at a strength of less than or equal to approximately 0.2 T and greater than or equal to approximately 0.1 T.
 22. The portable magnetic resonance imaging system of claim 1, wherein the B₀ magnet is configured to produce a B₀ field at a strength of less than or equal to approximately 0.1 T and greater than or equal to approximately 50 mT.
 23. The portable magnetic resonance imaging system of claim 1, wherein the B₀ magnet is configured to produce a B₀ field at a strength of less than or equal to approximately 50 mT and greater than or equal to approximately 20 mT.
 24. The portable magnetic resonance imaging system of claim 1, wherein the B₀ magnet is configured to produce a B₀ field at a strength of less than or equal to approximately 20 mT and greater than or equal to approximately 10 mT.
 25. The portable magnetic resonance imaging system of claim 1, wherein the electromagnetic shielding comprises at least one conductive strip connecting an upper portion and a lower portion of the magnetic resonance imaging system to form at least one conductive loop.
 26. The portable magnetic resonance imaging system of claim 25, wherein the portable magnetic resonance imaging system further comprises at least one ferromagnetic component configured to capture and direct at least some of the magnetic field generated by the B₀ magnet to increase the magnetic flux density within the imaging region of the magnetic resonance imaging system, and wherein the at least one conductive strip is coupled to the at least one ferromagnetic component to form the at least one conductive loop.
 27. The portable magnetic resonance imaging system of claim 26, wherein the B₀ magnet comprises at least one first magnet and at least one second magnet positioned in a bi-planar arrangement about the imaging region, and wherein the at least one ferromagnetic component comprises a ferromagnetic structure magnetically coupled to the at least one first magnet and to the at least one second magnet, and wherein the at least one conductive strip is coupled to the ferromagnetic structure to form a conductive loop about at least a portion of the imaging region.
 28. The portable magnetic resonance imaging system of claim 25, further comprising a plurality of fastening mechanisms configured to allow the at least one conductive strip to be removably attached at any of a plurality of different locations about the imaging region.
 29. The portable magnetic resonance imaging system of claim 25, wherein the at least one conductive strip is formed of a flexible material.
 30. The portable magnetic resonance imaging system of claim 29, wherein the at least one conductive strip is formed of a flexible braided material.
 31. The portable magnetic resonance imaging system of claim 25, wherein the at least one conductive strip is formed of a rigid material.
 32. A portable magnetic resonance imaging system comprising: a B₀ magnet configured to produce a B₀ field for an imaging region of the magnetic resonance imaging system; a noise reduction system configured to detect and suppress at least some electromagnetic noise in an operating environment of the portable magnetic resonance imaging system; and electromagnetic shielding for at least a portion of the portable magnetic resonance imaging system, the electromagnetic shielding providing substantially no shielding of the imaging region of the portable magnetic resonance imaging system. 33-46. (canceled)
 47. A portable magnetic resonance imaging system comprising: a B₀ magnet configured to produce a B₀ field for an imaging region of the magnetic resonance imaging system; a housing for the B₀ magnet; and at least one electromagnetic shield structure adjustably coupled to the housing to provide electromagnetic shielding for the imaging region in an amount that can be varied by adjusting the at least one electromagnetic shield structure about the imaging region. 48-75. (canceled) 